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Patent 2345632 Summary

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(12) Patent Application: (11) CA 2345632
(54) English Title: MIXED-MODE LIQUID VENTILATION GAS AND HEAT EXCHANGE
(54) French Title: ECHANGE DE GAZ ET DE CHALEUR AU MOYEN D'UNE VENTILATION MIXTE PAR LIQUIDE
Status: Dead
Bibliographic Data
(51) International Patent Classification (IPC):
  • A61F 7/12 (2006.01)
  • A01N 1/02 (2006.01)
  • A61M 16/00 (2006.01)
  • A61H 31/00 (2006.01)
(72) Inventors :
  • HARRIS, STEVEN BRADLEY (United States of America)
  • DARWIN, MICHAEL GREGORY (United States of America)
  • RUSSELL, SANDRA RENEE (United States of America)
(73) Owners :
  • CRITICAL CARE RESEARCH, INC. (United States of America)
(71) Applicants :
  • CRITICAL CARE RESEARCH, INC. (United States of America)
(74) Agent: SMART & BIGGAR
(74) Associate agent:
(45) Issued:
(86) PCT Filing Date: 1999-10-01
(87) Open to Public Inspection: 2000-04-06
Examination requested: 2004-09-23
Availability of licence: N/A
(25) Language of filing: English

Patent Cooperation Treaty (PCT): Yes
(86) PCT Filing Number: PCT/US1999/022833
(87) International Publication Number: WO2000/018459
(85) National Entry: 2001-03-26

(30) Application Priority Data:
Application No. Country/Territory Date
60/102,593 United States of America 1998-10-01

Abstracts

English Abstract




A technique, Mixed-mode Liquid Ventilation (MMLV), for ventilation, and heat
exchange is disclosed. The technique uses an endotracheal catheter (Figs. 1a
and 6b) to add, to remove liquid from the lungs of user
continuously/cyclically, and delivers gas at a rate independent of the
delivery of liquid. This technique produces small scale mixing of gas, liquid
in the user's airways, allowing efficient gas, and heat exchange. Medical uses
for the technique are disclosed, that include induction, reversal of
hyperthermia, and hypothermia.


French Abstract

L'invention concerne un nouveau procédé de ventilation et d'échange de chaleur appelé ventilation par liquide en mode mixte (MMLV : mixed mode liquid ventilation). Ce procédé comprend l'utilisation d'un cathéter endothrachéal (fig. 1a et 6b) permettant d'ajouter et de retirer du liquide des poumons de manière continue et/ou cyclique, et d'administrer un gaz à un débit indépendant de l'administration de liquide. Ce procédé produit un mélange de gaz et de liquide à petite échelle dans les voies respiratoires, permettant un échange efficace de gaz et de chaleur. L'invention concerne également des applications médicales de ce procédé, parmi lesquelles l'induction et l'inversion de l'hyperthermie et de l'hypothermie.

Claims

Note: Claims are shown in the official language in which they were submitted.




What is claimed is:

1. A method for gas and/or heat exchange comprising:

addition and removal of oxygenated liquid continuously and or cyclically
from the lungs and
delivery of gas at a rate independent of the input and removal of oxygenated
liquid allowing mixing of said gas and oxygenated liquid producing small scale
convection and
cooling or heating of said oxygenated liquid to a temperature in the range of
about -10 to about 43°C.

2. The method for gas and/or heat exchange according to Claim 1, wherein the
cooling or heating of said oxygenated liquid is to a temperature in the range
of about -10 to
about 20°C.

3. The method for gas and/or heat exchange according to Claim 1, wherein the
cooling or heating of said oxygenated liquid is to a temperature in the range
of about -10 to
about 0°C.

4. The method according to Claim 1 wherein said oxygenated liquid comprises
perfluorocarbon.

5. The method according to Claim 1 wherein said small scale convection
occurs during TLV or PLV.

6. The method of Claim 1 further comprising the addition of a technique
selected from the group consisting of: HFV, HFOV, Sweep flow and mixtures
thereof.

7. The method for gas and/or heat exchange according to Claim 1, wherein
nitric oxide or nitric oxide donors are administered to facilitate gas and
heat exchange.

8. The method for gas and/or heat exchange wherein the gas is helium.


9. A method for the treatment of hypothermic pathologies in a mammal
comprising the method of Claim 1 wherein said oxygenated liquid has a
temperature of
about 40°C to about 45°C, and said method is continued until the
temperature of said
mammal is between about 35°C and 39°C.

10. The method according to Claim 9 wherein the hypothermic medical
conditions are environmental exposure or cold water drowning.



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11. A method for the treatment of hyperthermic pathologies in a mammal
comprising the method of Claim 1 wherein said oxygenated liquid has a
temperature of
about -5°C to 30°C, and said method is continued until said
temperature of said mammal is
between 35°C and 39°C.

12. The method according to Claim 11 wherein the hyperthermic pathologies are
selected from the group consisting of heatstroke, malignant hyperthermia,
hyperpyrexia,
stroke, head injury, post-ischemic insult, and febrile illness.

13. A method for facilitating gas exchange more efficiently for the treatment
of
respiratory diseases in a mammal, comprising the method of Claim 1.

14. The method of Claim 13 wherein said respiratory diseases are selected from
the group consisting of adult and neonatal respiratory distress syndrome,
pulmonary edema,
alveolar proteinosis, chronic bronchiectasis, and chemical and thermal insults
to the lung.

15. A method for the preservation of biological material comprising the method
of
Claim 1.

16. The method of Claim 15 wherein said biological material is selected from
the group consisting of cells, tissues, organs, and bodies.

17. A method for increasing the efficiency of closed chest CPR comprising the
method of Claim 1 wherein the liquid loading is synchronized with chest
compression.

18. An apparatus for gas and/or hear exchange comprising:
a source of biocompatible oxygenated liquid,
a conduit having a first end connected to said source of oxygenated liquid and
delivered via two canulae directed down the main bronchi
a source of breathing gas;and
a computer to control the loading and unloading of said oxygenated liquid
continuously or intermittently,
a heat exchange/filter/oxygenating assembly, and
delivered in such a way as to induce convective mixing of said oxygenated
liquid
and said gas in said lungs.

19. The apparatus according to Claim 18 wherein said computer delivers liquid
at a liquid minute volume 25-50% of that for gas.

20. The apparatus according to Claim 18 furthur comprising a cold reservoir
for
cooling said oxygenated liquid.



-57-


21. The apparatus according to Claim 18 furthur comprising a heat exchanger
for heating said oxygenated liquid.

22. The apparatus according to Claim 18 wherein said computer monitors
infusion and suction pressures to minimize them subject to the amount of
liquid necessary
for heat exchange and the amount of CO2 and O2 which need to be removed and
deleted
respectively in gas exchange.

23. The apparatus according to Claim 18 wherein said computer monitors liquid
infusion and removal temperatures in order to maximize heat exchange subject
to the
amount of liquid necessary for heat exchange and the amount of CO2 and O2
which need to
be removed and deleted respectively in gas exchange.

24. The apparatus according to Claim 23 wherein said liquid infusion and
removal temperatures are selected from the group consisting of: venous,
arterial, and both
venous and arterial temperatures.



-58-

Description

Note: Descriptions are shown in the official language in which they were submitted.



CA 02345632 2001-03-26
WO 00/18459 PCT/US99/22833
Mixed-Mode Liquid
Ventilation Gas and Heat Exchange
FIELD OF THE INVENTION
The present invention relates to ventilator and heat exchange systems and,
more
particularly, to a "mixed-mode" gas-plus-liquid ventilator system using an
endotracheal
catheter to add and remove liquid ventilation or heat-exchange medium from the
lungs
continuously and/or cyclically, with delivery of gas to the lungs at a rate
and volume
independent of addition and removal of liquid.
BACKGROUND OF THE INVENTION
There are many situations in both human and veterinary medicine where it is
desirable to rapidly induce or reverse hypothermia. There are also many
clinical situations
where it is essential to be able to rapidly reduce dangerously elevated body
temperatures to
near normal, as in the case of hyperthermia from heat stroke, drug or surgical
anesthetic
reaction, and febrile illness secondary to stroke, infection or other
illnesses. In fact, it has
been demonstrated in a number of studies that patient mortality is directly
dependent on
high temperature exposure time, and inversely dependent on the rapidity with
which core
temperature is normalized.
Heretofore, the only clinically available means of achieving very rapid
reduction in
body temperature (or conversely, of re-warming from hypothermic temperatures)
has been
the use of invasive methods of heat exchange, such as cardiopulmonary bypass
(circulating
blood over a heat exchanger), or peritoneal and/or pleural lavage. A third,
slower
alternative for changing body temperature involves immersing the patient in a
bath of
heated or chilled liquid or gas (e.g. helium). The problems with these
approaches are many:
1 ) External means of chilling or re-warming are relatively slow (< 0.01
°C to 0.20°C/min),
and produce a host of undesirable and sometimes lethal complications. In the
case of
cooling, the chilling of external body tissues results in vasoconstriction
which interferes
with the delivery and removal of oxygen, nutrients, and wastes from the
peripheral tissues.
2) Re-warming from hypothermia 1 by external means can cause a peripheral vaso-

relaxation, hypotension, and effective hypovolemia, for which the cold-
impaired heart and
autonomic nervous system cannot compensate. Profound hypotension may develop,
causing cardiac arrest and death, sometimes paradoxically in people presenting
for
apparently non-critical conditions.


CA 02345632 2001-03-26
WO 00/18459 PCT/US99/22833
3) Peripheral tissues being re-warmed recover the need for oxygen and
metabolic substrates
before the circulatory system and other organs can deliver them (since these
organs are still
cold and functioning marginally). This resulting imbalance between metabolic
supply and
demand results in the generation of large amounts of anaerobic waste products,
including
carbon dioxide and lactate, which decrease blood and tissue pH and result in
severely
disturbed homeostasis.
4) If external re-warming proceeds without inducing cardiac arrest, a second
phase of risk
occurs when "after-drop" is experienced. After-drop is a reduction in body
core
temperature during slow external re-warming. After-drop occurs as a result of
peripheral
vasodilation during patient re-warming, thus allowing large amounts of blood
to flow
through deeply chilled peripheral tissues, resulting in a seemingly
paradoxical drop in body
core temperature. After-drop can result in cardiac arrest during patient re-
warming if the
heart is cooled below its critical threshold for fibrillation. Though some
controversy exists
about the relative importance of this process in humans, it still remains of
great concern to
specialists in the field.
5) The use of invasive temperature modifying techniques such as peritoneal and
pleural
lavage, extracorporeal perfusion, or central venous cooling, are either not
very effective
(e.g. lavage techniques), or can be performed only in a medical setting by
highly skilled,
licensed practitioners (e.g. physicians). Most importantly, these techniques
cannot be safely
or reliably performed in the field by paramedics or other non-physician
emergency medical
personnel. In the case of techniques which require vascular access, many
medical facilities
possess neither the complex and costly equipment required to carry out such
procedures,
nor the highly skilled personnel necessary to perform such procedures. A
particular
problem with these methods is the need for bulky, complicated, failure-prone
equipment
which may be difficult to store in states of readiness (e.g. cardiopulmonary
bypass
apparatus). Technical errors and mechanical failures associated with
extracorporeal
techniques carry a high risk of morbidity, with such errors frequently
resulting in
neurological damage or loss of life.
The Lungs As a Gas Exchanger and Heat Exchanger
An alternative to invasive temperature modifying techniques would be to use
the
large surface area of the lungs as a heat exchanger. Nearly all of the cardiac
output (i.e., all
blood flowing to the body) flows through the lungs, and since the lungs
possess a surface
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CA 02345632 2001-03-26
WO 00/18459 PCT/US99/22833
area of at least 70 square meters, they form an ideal heat exchanger that
would allow for
rapid core cooling and re-warming of the patient without the problems
associated with the
techniques previously discussed. In addition, since the lungs are accessible
via the trachea,
the relatively benign maneuver of endotracheal intubation (a skill universally
possessed by
paramedics) allows for quick field access to this powerful heat exchanger.
The potential utility of the lungs as a heat exchanger was first recognized by
Clark and
Gollan in the 1960's, when they used the perfluorochemical FX-80 to
demonstrate the
concept of total liquid breathing in mice. The concept of using the lungs as a
heat
exchanger for therapeutic purposes was first proposed by Shaffer et al. in
1984, using total
liquid ventilation and the fluorocarbon "Rimar 101" (Rimar Chimica S.p.A.,
Vincenza,
Italy).
Heat exchange in the lungs using liquid ventilation is superior to gas
ventilation
because at standard temperature and pressure, gases such as oxygen and air
have only
approximately a 2200th of the volumetric specific heat capacity of water.
Thus, under
ordinary circumstances the lungs serve as a relatively poor heat exchanger if
only gaseous
media are used. This includes the use of the highly conductive-low viscosity
gas mixture
of oxygen and helium (Heliox). The high conductivity of Heliox makes it far
more
efficacious as a heat exchange medium under high pressure conditions where its
specific
heat capacity is greater that at normal pressures; however these conditions
are of little
relevance to most clinical situations.
The Basics of Liquid Ventilation
Liquid ventilation involves the breathing of gas-carrying liquid as the medium
of
gas exchange within the lungs. Since the first liquid ventilation experiments
(1950's) in
mice using super-oxygenated saline, several liquid media for ventilation have
been studied.
The class of agents currently optimized to function as liquid breathing media
are the
fluorocarbons (containing only fluorine and carbon), and the organic
perfluorochemicals
(PFCs). PFC compounds contain elements other than fluorine and carbon, with
fluorine or
other halogens comprising the majority of peripheral moieties within the
molecule. As a
class, PFC compounds comprise molecules that are relatively insoluble in
either water or
lipid, and are more-or-less chemically and pharmacologically inert. PFCs do
not dissolve
native lung surfactants, and are far less injurious to the lungs than any
known silicone or
water-based solution.
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CA 02345632 2001-03-26
WO 00/18459 PCTNS99/22833
Total Liquid Ventilation
Historically, the first mode of liquid ventilation studied was total liquid
ventilation
(TLV). In TLV all of the gas within an animal's lungs is replaced with liquid,
and each
breathing cycle (tidal volume) is composed entirely of liquid medium. While
this modality
holds promise in deep-sea diving research, it has not yet been used in humans.
In heat exchange, each TLV breathing cycle provides a certain volume that can
be
passed through the lung heat exchanger. As noted, the advantage of TLV (as
opposed to gas
exchange) for heat exchange, is that liquids such as PFCs have a specific heat
capacity
several thousand-fold that of gas at normal pressure. Despite this advantage,
total liquid
ventilation suffers from a number of drawbacks:
1) In TLV it is necessary to completely eliminate air from the animal's lungs
and the
ventilating circuit, because PFCs do not pump well in many systems due to
"vapor-lock".
The maneuvers necessary to clear all gas from the system are problematic and
time
consuming.
2) Due to the increased viscosity of liquids (PFCs are 80 times more viscous
than
air), the number of liquid breaths attainable per minute is sharply
constrained compared to
air ventilation. Typically, no more than 5 to 7 liquid breaths per minute are
possible
(Shaffer TH, et al., 1984). This is approximately one forth the usual rate at
which tidal gas
volume ventilation occurs in animals of this size.
3) In addition, the maximal liquid ventilatory "minute volume" (dV/dt) is more
tightly constrained in TLV making adequate gas exchange problematic in
situations where
oxygen demand is high, and the need to remove CO~ is great. Carbon dioxide
removal is a
particular problem in TLV because PFCs have a lower carrying capacity for COZ
at
physiologic partial pressures (which cannot be changed much), than they have
for 02
(which is easily deliverable at artificially high partial pressures).
Miyamoto and Mikami in 1976 calculated that the resting man produces normally
192 mL/min of COZ (S.T.P.). This level of COz production would require TLV
(PFC)
ventilation volumes of about 4 L/min (or about 70 mL/kg/min). Although this is
only 70%
of the normal gas ventilatory flow for a resting adult, it is near the upper
limit of flows that
can be accomplished at normal pressures in TLV (Kylstra, 1974). The higher
peak and
mean ventilating pressures necessary to move the amount of liquid required for
CO~
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CA 02345632 2001-03-26
WO 00/18459 PCTNS99/22833
exchange in TLV would expose the lungs to an increased risk of baro-trauma
(pressure
injury) and volu-trauma (over-distention injury).
During higher than normal COZ production rates (e.g. disease), TLV would
clearly
not be adequate for COz removal. Examples of high CO, (hypercapnic) states are
1)
increased metabolic states (e.g. cancer, infection, burns), 2) states of
physiologic stress (e.g.
hyperthermia, agitation), and 3) post-ischemic conditions where substantial
metabolic debt
has been incurred and the need to rapidly unload CO~ and deliver large amounts
of OZ are
essential. Such hypercapnic/hypercarbic states are also frequently present in
shock due to
sepsis or trauma, and thought to be due to both an increased production of
CO2, and a
decreased elimination of COZ due to low blood flow or pulmonary edema.
In anesthetized, paralyzed, normothermic dogs, TLV is capable of maintaining
steady-state gas exchange with adequate O, delivery and CO, removal. However,
TLV is
not adequate to steady-state CO, removal under basal metabolic conditions in
smaller
animals with higher specific metabolic rates, such as guinea pigs. As Matthews
and co-
workers document ( 1978), the parameters for maintaining normocapnia in
anesthetized
beagles are narrow, even under basal normothermic metabolic conditions. In
this study, as
liquid ventilation rates were increased from 2.8 to 5.6 liquid breaths per
minute, and
alveolar ventilation was increased from 574 to 600 mL/min/animal (increase of
4%), the
paCO~ continued to increase until dangerous hypercapnia occurred. The authors
suggested
that this increase was due to a 2% drop in liquid-alveolar ventilation,
however using their
own formulas and data, we have calculated that the dogs receiving higher
ventilation rates
actually have higher rates of alveolar ventilation (dVa/dt). These results
would seem
paradoxical until consideration is given to the inverse relationship of paCO,
to alveolar
ventilation, a relationship which holds only if equilibrium between blood and
"alveolar"
C02 (actually, alveolar and small airway C02) is reached for each breath. The
fact that
high TLV ventilatory rates resulted in rising paCO, in this paper despite
increased
"alveolar" ventilation (dVa/dt), indicates that the high ventilatory rate used
was too rapid
for blood/airway C02 equilibrium to be reached. This behavior is a limitation
of the
diffusion speed of CO, in PFC in the small airways, conduits that are
constructed on a size
scale for "gas-in-gas" diffusion, but not "gas-in-liquid" diffusion.
Diffusion limitations on CO, removal in TLV models have been noted by several
research groups. This diffusion limited failure of C02 equilibrium acts to
increase
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CA 02345632 2001-03-26
WO 00/18459 PCTNS99122833
"diffusion dead space" and in practice places even more constrictive limits on
usable
ventilation rates with perfluorocarbon (PFC) liquids in TLV. These limits are
in addition to
those already imposed by the viscosity of the fluid itself. For example, Koen
and Shaffer
found that TLV in young cats showed maximal CO, elimination at a ventilatory
rate of 3 to
3.5 breaths per minute. Decreasing CO, clearance occurred at lower rates, due
to
insufficient ventilation, whereas decreased CO, clearance occurred at higher
rates due to
COZ diffusion limitations with short liquid dwell times. In summary, at higher
TLV
ventilatory rates, equilibrium in pC02 between alveolar blood and freshly
inspired liquid in
the small airways is not reached. For this reason, liquid alveolar ventilation
in TLV cannot
be arbirarily increased, for fundamental reasons involving pressures limits on
high liquid
flows, and also diminishing gas exchange at rapid liquid flow rates.
4) PFC viscosity (pressure/flow) also places a limit on the rate at which heat
can be
extracted from an animal or patient using TLV. In addition to the CO,
diffusion limitation,
there is indirect evidence suggesting that thermal equilibrium is not reached
between blood
and liquid in small airways at high TLV "alveolar ventilation" rates. Thus,
there appears to
be a heat-diffusion limitation to TLV that is analogous to the CO, diffusion
limitation.
This phenomenon may explain why Shaffer's TLV cat studies failed to achieve
concomitant increases in the rates of animal core cooling, when significantly
greater PFC
temperature gradients were used (Shaffer TH, et al., 1984). In Shaffer's
report, it was found
that decreasing PFC infusion temperature from approximately 20°C to
about 10°C (from 0
T = 1 S°C to 0T = 24°C), resulted in cooling rates increasing
from 0.13°C/min (7.8°C/hr) to
0.15°C/min (9.0°C/hr), a change of only 15%. This 15% increase
occurred despite an
increase of ~T equal to 60%. These results suggest a sharp decline in the
efficiency of heat
extraction with increased 0T at higher TLV ventilation rates (in this
experiment, rate was
increased from 4.5 to 5.3 liquid breaths/min).
In Shaffer's study, the authors calculate from PFC inspiration and expiration
temperature differences, a 96% increase in heat extraction per kg from their
animals at the
10°C PFC infusion temperature versus that calculated at 20°C.
However, the fact is that this
increase in heat extraction does not show up in the rate of body core cooling
(15%), to
which it should be proportional. This indicates that Shaffer's calculations of
heat removal
performed on the basis of integrated measurements of expired fluid
temperatures must have
been in error. As further evidence of this error, calculations of expected
cooling rates of
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CA 02345632 2001-03-26
WO 00/18459 PCT/US99/22833
animals used in this study (using a reasonable 0.8 cal/g/°C. or
kcal/kg/°C average specific
heat capacity for the body), indicate that up to half of the heat extraction
calculated by PFC
temperature differences in this experiment are unaccounted for even at the
fastest cooling
rates. For example, an animal with an average 0.8 kcal/kg/°C specific
heat capacity, cooling
S at the reported rate of 9.0°C/hr, could theoretically give up heat at
a rate no faster than (0.8
kcal/kg/C)(4184 J/kcal)(9 C/hr) = 30,124 J/kg/hr. However, Shaffer's
experiment reports
on the basis of temperature readings of PFC infused and expired, the
extraction of 65,637
J/kg/hr. It is likely that the difficult integration of [expired,fluid
temperature] versus [fluid
volume] curve for this experiment was in error by a factor of 2Ø For
examples of
experiments in which integrated cooling rates calculated from PFC temperature
differences
match actual animal body cooling, see the canine experiments using hand
controlled
infusion below. The authors of the present patent have found that at rapid
(machine-
controlled) liquid infusion and removal rates, peak fluid temperatures do not
accurately
reflect volume-averaged fluid temperatures, or fluid heats.
Partial Liquid Ventilation
The second mode of liquid ventilation to be studied was Partial Liquid
Ventilation
(PLV). In PLV, the subject's lungs are partially (usually to functional
residual capacity
(FRC) of 30 mL/kg of body weight or about 1/3rd of total lung capacity
(assumed hereafter
to be 90 mL/kg) loaded with PFC liquid. In PLV, PFC liquid loading is
accompanied by
conventional mechanical ventilation using a standard gas ventilator at normal
gas rates and
tidal volumes. Since the breaths are delivered as gas, PLV allows for the
number of breaths
per minute, and alveolar ventilation rates, to be set much closer to the
physiologically
acceptable and desirable rate. PLV can even be used with high frequency gas
ventilators,
and can accommodate a wide range of metabolic states in which the demand for
02
delivery and C02 removal is greater than that of basal states. PLV is
currently being tested
in human clinical trials.
During PLV, gas exchange occurs across material boundaries at two locations:
1)
between the PFC liquid and the circulating blood (across the alveolar
membrane), and 2)
between the PFC liquid and the ventilating gas in the airways, where a short-
lived turbulent
foam of PFC and ventilating gas is created. The low viscosity of this PFC foam
allows it to
reach briefly into the small airways of the bronchial tree with each breath
and helps explain
the complex and poorly understood mechanism of PLV gas exchange. The turbulent
mixing
_7_


CA 02345632 2001-03-26
WO 00/18459 PCT/US99/22833
of PFC foam may also explain the newly appreciated heat exchange properties of
this
modality (as well as those of Mixed-Mode Liquid Ventilation, discussed next).
The mixing
of air and gas in small airways, which will be discussed more fully later,
appears to be key
to improving the heat transfer limitations of TLV. We believe that the mixing
of PFC and
gas disrupts laminar liquid (PFC) flow in small airways by introducing
turbulence to the
fluid, thereby improving the small-scale (small airway) convection necessary
for maximal
heat transfer rates.
We introduce in this document the novel application of the unique mixing
features
of PLV to assist in core body-heat transfer. Specifically, the primary
clinical utility of PLV
has heretofore been in the treatment of adult and neonatal respiratory
distress syndromes.
In these pathological conditions, PFC (among other salutary effects) moves
down ward in
the bronchial tree due to its high density ( 1.8 to 2.0 times that of water)
opening alveoli
which are closed as the result of pulmonary edema {fluid in the dependent
portions of the
lungs).
Significant heat transfer has not been documented using standard PLV because
PFC
has been historically loaded slowly into the lungs, and once in place, has not
been retrieved
(cycled). Using a single dose of 2/3rds or more of total lung capacity (TLC)
of cold
(slightly below 0°C) PFC (60 mL/kg), it is possible in the dog to
achieve a uniform core
cooling of approximately 1.5°C with only modest injury from baro- and
volu-trauma {see
Example 1 data in Part II below). Further cooling of the test subject does not
occur unless a
new load of cold PFC is instilled into the lungs. Thus, the use of PLV and
single PFC
loads, even to extreme volumes (i.e., those approaching TLC) is not a viable
means for
achieving even moderate, controllable, or lasting hypothermia.
Note On the Difficulty of Re-warming
Single loads of PFC are not sufficient to induce significant or lasting
hypothermia.
This problem is compounded more so when contemplating PLV for hypothermic
subject
re-warming. This is because there are two unavoidable limitations on how warm
the
delivered liquid can be. The first is that an absolute temperature limit of
42°C exists beyond
which hemolysis and acute thermal injury to tissue occurs. The second limit
involves the
temperature gradient (DT) between the blood and the PFC liquid, which if
significantly
greater than 5°C, exposes the subject to the risk of gas bubble emboli.
This risk occurs
because the solubility of nitrogen and other gases in plasma is greater at
cold temperatures.
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CA 02345632 2001-03-26
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Upon re-warming chilled blood, the nitrogen (and other gases 1 come out of
solution,
forming gas bubbles which can then embolize both the arterial and venous
circulation. This
same phenomenon occurs in nitrogen saturated tissues warmed at very rapid
rates, and
forms the pathological equivalent of the "bends" experienced by deep sea
divers breathing
nitrogen-containing gas as they decompress too rapidly.
These temperature gradient constraints on warm liquid delivery to hypothermic
subjects sharply limits the maximal therapeutic rate of heat transfer
achievable by PLV
used for re-warming. For example, it is not possible to deliver liquid that is
35°C above
body temperature in a subject still warm enough to have a beating heart
(typically 25°C or
above). In contrast, it is possible (see Part II) to safely deliver liquids
that are 35°C cooler
than body temperature. Thus, the OT when re-warming from modest but life-
threatening
hypothermia (i.e., 27°C) is less than 30% of that which can be achieved
during the
therapeutic induction of hypothermia.
SUMMARY OF THE INVENTION
One aspect of the present invention is a mixed-mode liquid ventilation (MMLV)
method for gas and/or heat exchange in the lungs (human clinical and
veterinary
applications). The MMLV method allows mixing of gas and liquid in the small
airways
of the lungs, producing small-scale liquid mixing in a convection-like
process, rapid
return of fluid from the lung periphery, and more rapid and efficient transfer
of heat, and
dissolved gasses, during the practice of ventilation with liquids.
In one embodiment, nitric oxide or nitric oxide donors are administered to
facilitate
gas and heat exchange.
In another embodiment the gas is helium.
In a further embodiment the liquid ventilation medium is a perfluorocarbon or
perfluorochemical.
A further aspect of the present invention is a method of inducing small-scale
mixing
of liquid heat exchange ventilation media, using other known ventilation
methodologies,
alone or in combination. These specifically include known types of gas
ventilation,
including high frequency oscillating ventilation.
Another aspect of the invention is the use of MMLV to treat hypothermic
pathologies by heating said liquid ventilation medium, and thus increasing
body
temperature.
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A further aspect of the invention is the use of MMLV to induce hypothermia for
medical purposes, or to treat hyperthermic pathologies, by cooling said liquid
ventilation
medium, and thus decreasing body temperature. In this embodiment, liquid
ventilation
media may be infused at temperatures as low as -10 C. This is made possible by
the
presence of thermally buffering PFC already in the lung, as well as the fact
that PFC may
be a minor volumetric component of the ventilation mix in MMLV, and small
infusions
are warmed to temperatures above 0 C before freezing or chilling damage can be
done to
tissues.
Another aspect of the invention is a method of preserving biological material,
for
example beating-heart cadaveric preparations, using the rapid cooling
available witth
MMLV .
A further aspect of the invention is a method for increasing the efficiency of
CPR
using MMLV.
A final aspect of the invention is an automatic apparatus for MMLV, which
connects to the lungs via the bronchi and uses a computer to control loading
and
unloading of said oxygenated liquid and gas so that mixing occurs; and also
makes use of
the computer to insure that pressure limits are not exceeded, and gas
ventilation proceeds
in a way which most rapidly induces removal of fluid heat exchange media. In
this
embodiment, the computer controls liquid infusion in such a way as to maximize
time
integrated arterial/venous temperature differences, for best cooling rates,
subject to
ventilatory constraints.
In a typical embodiment said computer controls the liquid ventilatory volume
delivered and removed (dV/dt) to be about 10% to 50% that typically necessary
for gas
ventilation.
In a further embodiment the apparatus has a cold reservoir where pre-cooled
PFC
may be stored.
In another embodiment the apparatus has a heat exchanger for PFC.
In another embodiment the apparatus has an active liquid ventilation system,
which
may employ a canula able to remove liquid at the same time gas breaths are
delivered.
Gas and liquid are typically infused and removed through separate concentric
tubes in
many of the most efficient implimentations of the invention, but may be
removed through
the same tube.
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Further objects, features, and ether aspects of the present invention become
apparent from the ensuing detailed description, considered together with the
appended
figures.
BRIEF DESCRIPTION OF THE FIGURES
FIG. 1 shows a graph from Example 1 of canine tympanic, central venous blood,
aortic arterial blood, and esophageal surface temperatures over time, during
mixed gas and
liquid ventilation, in which a single loading volume of liquid PFC is
delivered into the
endotracheal tube. PFC as a single load of 76 mL/kg of PFC at a temperature of
4.4°C is
loaded into the lungs over 145 seconds.
FIG. 1 a shows a canine instrumented for measurement of temperatures as in
Example
1. Indirect brain temperature is measured via a copper/Constantan thermocouple
probe
placed on the right tympanic membrane. Venous temperature was measured by a
thermistor in a thermodilution catheter inserted via the femoral vein into the
inferior vena
cava. Arterial blood temperature was measured by a thermocouple probe inserted
via the
femoral artery into the descending aorta. Selected instrumentation modalities
from
Examples to follow are also illustrated.
FIG. 2 illustrates the three major thermal kinetic compartments, and
respective
fractional heat capacities of such compartments, in an anesthetized canine
with normal
cardiac output, undergoing heat exchange through PFC lung lavage, as in
Example #1.
Existence and characteristics of such compartments may be inferred from
measurements
such as are graphed in FIG. 1, in studies such as Example #I, in the way that
is hereafter
detailed.
FIG. 3 shows canine tympanic, rectal, venous blood, and aortic blood
temperature
as a function of time from Example 2, showing multiple cycles of loading and
unloading of
chilled PFC at -1°C from the lungs. This example illustrates PFC
removal using a suction
reservoir, illustrating the full principle of mixed-mode liquid ventilation
(MMLV). A 25.7
kg anesthetised. intubated, and paralyzed dog was given infusions of PFC at a
maximal rate
of 6 mL/kg/min. Temperatures were measured as in FIG. 1 a.
FIG. 4 is shows a table of PFC liquid infusion volumes, measured temperatures
of
these volumes, and the calculated difference in their heat contents, from
example 2. This
table illustrates the qualitative and quantitative heat transfers involved in
the MMLV
technique.
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FIG. 5 shows a graph of canine tympanic, central venous. aortic, rectal, and
PFC
suction temperatures from Example 3 of a 24.5 kg dog when infusion and suction
rates are
increased to 16.7 mL/kg/min. Temperatures were measured as in FIG. 1 a.
FIG. 6A is an illustration of a device for Mixed-mode Liquid Ventilation
(MMLV).
Its use is in manual mode (manual valve control) is illustrated in Examples 2
through 6. Its
use in full computer-valve-control mode is illustrated in figures I Oa and 19.
FIG.6B is an illustration of how the device of FIG.6A attaches to the lungs
via the
endotracheal tube.
FIG.7 shows canine tympanic, rectal, venous, and aortic blood temperature as a
function of time from Example 4. A 17.3 kg dog was used with an infusion rate
of 45
mL/kg/min, increased by a factor of 2.6 from that shown in FIG. 3. Illustrated
is a
procedure which cooled the animal by 12 C over 30 minutes. for net cooling of
10 C after
thermal equilibration. This animal survived long-term, without evidence of
respiratory
damage. Temperatures were measured as in FIG. 1 a.
i 5 FIG.B shows canine rectal temperature as a function of time from Example 6
of a 30
kg dog (a much larger animal than illustrated in FIG.7), when the infusion
rate was
maintained at 46 mL/kg/min, and suction rates were increased to accommodate
larger
absolute volumes. Temperatures were measured as in FIG. la.
FIG. 9 shows rectal, tympanic, venous, aortic, and suction temperature as a
function
of time from Example 6. using a 19.8 kg dog with PFC infusion rate of 46
mL/kg/min.
This example illustrates temperature relationships during the fastest manual
cooling cycle
rates. Temperatures were measured as in FIG. la.
FIG. 9a shows the same experiment, but with temperatures graphed without PFC
temperatures, for ease of interpretation.
FIG. 10 Shows typmanic temperatures of 5 anmals cooled by rapid lavage with
PFC by MMLV techniques, vs Surface cooling.
FIG. l0a shows tympanic temperatures of 5 animals cooled by rapid lavage (same
group as above) vs cooling of dogs by machine cooling at 50% of the lavage
load, but twice
the lavage rate.
FIG.11 a shows P02 results of arterial blood gases drawn every 2 mintues in 4
animals cooled by rapid manually controlled PFC lavage, vs 2 animals given the
same
manual lavage with body temperature PFC. Temperatures are measured as in FIG
la.
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FIG.11 b shows p02 results of arterial blood gases drawn every 2 mintues in 3
animals cooled by rapid machine controlled PFC lavage, vs 2 animals given the
same
maschine lavage with body temperature PFC. Temperatures are measured as in FIG
la.
FIG.11 c shows paC02 results of arterial blood gases drawn every 2 mintues in
5
animals cooled by rapid manually controlled PFC lavage, vs 2 animals given the
same
manual lavage with body temperature PFC. Temperatures are measured as in FIG 1
a.
FIG.lId shows paC02 results of arterial blood gases drawn every 2 mintues in 3
animals cooled by rapid machine-controlled PFC lavage, vs 2 animals given the
same
lavage lavage with body temperature PFC. Temperatures are measured as in FIG
la.
FIG.12 shows end-tidal carbon dioxide concentration (pCO,) and minute carbon
dioxide production (VCOZ = dV/dt) during a MMLV induced decrease in canine
body
temperature of 5°C. VC02 was measured using a Novametrix Medical System
COZSMO
Respiratory Profile Monitor inserted into the ventilator circuit above the
endotracheal tube.
FIG.13 Shows VC02 measured by the above method in another dog during another
rapid PFC manual MMLV lavage. This graph shows that VC02 production is stable
after
PFC lavage stops, and does not rise until shivering begins. The slight
increase immediately
after lavage is known to be false spectrophotometric reading of PFC vapor.
FIG. l ~ shows the relationship between aortic temperature and heart rate
during cold
MMLV lavage in a canine. Temperatures were measured as in Fig 1.
FIG.15 shows the central venous pressure and heart rate during MMLV. Central
venous pressure was monitored via a fiber optic pulmonary artery (PA) catheter
advanced
via open cut down of the femoral artery to the level of the right atrium.
FIG.16 shows the central venous pressure and mean arterial pressure during
MMLV. Arterial pressure was monitored via a line inserted into the abdominal
aorta via
open cut-down of the femoral artery. Central venous pressure was monitored as
in FIG. 15.
FIG 16a shows the same study and variables graphed for a different canine
being
lavaged in the same way with body temperature PFC, in order to eliminate the
temperature
variable.
FIG.17 shows the central venous pressure and PFC infusion/suction temperature
during MMLV. This illustrates the relationship between infusion of cold PFC
and venous
pressure. Central venous pressure was monitored as in FIG. 15.
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FIG 17a shows the central venous pressure and PFC infusion/suction temperature
during MMLV, as in FIG I7, but using infused PFC at body temperature.
FIG.18 shows the PFC infusion temperature and ventilator airway pressure
during
MMLV with cold PFC.
FIG.19 shows airway pressure over time with cold machine controlled rapid PFC
MMLV cycling, as in FIG.s 10a. Here, pressure control by machine is especially
successful.
FIG 20 shows airway pressure over time with cold manual PFC MMLV cycling, in a
canine from FIG 10. Here, pressure control by hand is especially good.
DETAILED DESCRIPTION OF THE INVENTION
We have developed a new method for gas and heat exchange termed Mixed-Mode
Liquid Ventilation. Additionally, we have developed an apparatus which allows
for
heating and cooling by lung lavage while causing the least damage during the
process.
The technique and apparatus also may be used with liquid ventilation to
increase C02
1 S removal from the body, when liquid media have been introduced to the lungs
primarily
for heat exchange purposes. The technique and apparatus also take advantage of
our
discovery that gas ventilation may be used to facilitate recovery of liquid in
perfluorocarbon lung lavage, as well as the novel discovery that the non-
thermal
convection (transport of dissolved gas by liquid mass flow not related to
differential
density) that occurs with mixing of liquid and gas in the lungs' small
airways, allows for
faster and more efficient heat and gas exchange.
Mixed-Mode Liquid Ventilation
A technique is needed to achieve better and more efficient cooling or re-
warming
using liquid ventilation. This technique must allow for the continuous
addition and
removal of PFC (as liquid or aerosol) from the lungs, while also allowing for
the delivery
of gas breaths via a mechanical ventilator or other means at a rate
independent of the input
and removal of PFC liquid from the lungs is required. The reason for this is
that while
liquid ventilation is performed to control addition or removal of heat, gas
ventilation must
also occur independently to allow for adequate oxygen delivery and carbon
dioxide
removal. Furthermore, the combination of these techniques affords a previously
un-
described synergism to occur (as set forth herein), in that gas ventilation
assists liquid
removal from the lungs, and that gas-induced mixing of fluids in the small
airways is
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required to obtain optimum heat transfer efficiency during short liquid
ventilation cycle
times. In a new approach to PLV, our experiments suggest the possibility of
utilizing
some elements of PLV to facilitate heat exchange. We have discovered that when
optimized, PLV by virtue of PFC-gas bubble induced liquid convection,
possesses both
excellent heat and gas exchange properties.
That gas-bubble-induced PFC mixing (small-scale convection, not necessarily
thermally induced) must occur in PLV in order to obtain maximal efficiency of
heat
transfer in liquid ventilation, is not obvious. It is also not obvious that
independent control
of gas ventilation is necessary for the most rapid return and cycling of
liquid/gas mixtures.
A novel method of optimized gas and heat exchange with gas plus PFC liquid
ventilation is
detailed below. Over four years of work was required to develop the mechanics
involved in
"Mixed-Mode Liquid Ventilation" (MMLV) and to describe the general principles
by
which both heat exchange and gas exchange can proceed both simuhaneously and
independently from each other.
Our solution to the shortcomings of TLV and PLV (detailed below), was to
develop
a modality whereby PFC (which is appropriately chilled or warmed) could be
delivered to
the lungs, either continuously or intermittently, while conventional
mechanical ventilation
proceeded in a variable, but coordinated fashion. In the following discussion,
the specifics
of combined gas/PFC liquid ventilation and heat exchange, termed "Mixed-Mode
Liquid
Ventilation" will be outlined, and key experiments will be used to illustrate
the modality's
power. Finally, a prototype system for "field" MMLV application will be
revealed and
differentiated from prior art.
The Special Importance of PFC/Gas Mixing in the Periphery of the Lung
Before Mixed-Mode Liquid Ventilation (MMLV) is explained, further attention
should be drawn to a key phenomenon which is not taken advantage of in prior
art in liquid
breathing heat transfer (i.e., Total Liquid Ventilation/TLV heat transfer).
This phenomenon
occurs because PLV induces air and gas mixing in not only large airways, but
also in small
and extremely small airways.
Ordinary models of gas and liquid foams, based on high surface-tension
mixtures of
water and air, do not suggest the fineness of the gas micro-bubbles which
rapidly travel to
the extreme periphery of the lung in PLV. The radiopacity of the clinically
used
brominated PFC perflubron (LiquiventT"") has prevented this micro-bubble
formation
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phenomenon from being viewed directly by fluoroscopy. However, this phenomenon
may
be viewed directly by fluoroscopy if a non-radiopaque PFC (one which uses no
halogen
other than fluorine) is used. In such circumstances, each breath of PLV shows
a flash of
micro-bubbles rapidly (< 1 second) moving to the periphery of the lungs.
Liquid/gas
mixing and the resulting turbulence and disruption of insulating laminar flow
from such a
process is profound, and due to the unique fluid properties of PFCs, is
somewhat outside
ordinary experience. However, gas-induced "convective" mixing of PFC in small
airways
probably underlies not only the success of PLV in removing C02, but also the
increased
efficiency of PLV in heat transfer.
This relative efficiency has not been previously appreciated in PLV, partly
because
PLV, though now backed with significant clinical trial experience for
respiratory disease,
has not been used for heat transfer. As for TLV, it has not been used in
humans at all
(except briefly in a handful of test subjects), so that diffusion dead space
limitations
resulting from PFC in small airways is not yet encountered clinically. If
expected on
theoretical grounds in PLV, its absence may have been explained away in
clinical trials on
the grounds that many alveoli are not filled with PFC in this modality.
However,
radiographically the opposite appears to be the case, and especially in use of
PLV with high
frequency ventilation, it seems clear that essentially all peripheral airways
are filled with
PLV at all times, and yet TLV-type diffusion barriers still do not occur. Non-
thermal
"convective" small-scale liquid-mixing from high frequency oscillation induced
by such
ventilators may be the answer to a question which has not been asked.
The phenomenon of induced small-scale convection cannot be fully utilized for
heat
transfer efficiently unless PLV is combined with certain techniques for rapid
PFC liquid
infusion and removal, as detailed below, in the novel technique referred to
herein as Mixed-
Mode Liquid Ventilation. Mixed-Mode Liquid Ventilation (MMLV) is different
from TLV
and PLV, in being optimized for heat transfer.
Exposition of Mechanics of Mixed-Mode Liquid Ventilation
In Mixed-Mode Liquid Ventilation (MMLV), conventional mechanical ventilation
is initiated via endotracheal intubation or the use of other means to isolate
the airway from
the gastrointestinal system, and allow application of positive pressure to the
lungs (i.e., the
esophageal gastric tube obturator airway (EGTA), the Combitube airway, etc.).
Then,
chilled or warmed (relative to the subject's desired body temperature) PFC is
progressively
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loaded and unloaded from the lungs using a mechanical pump(s). The PFC may be
loaded
in liquid or aerosol/spray form, and loading may be done continuously or
intermittently
(timed with inhalations). The technique works best when the PFC is recycled
through a
gas-exchanger as well as a heat exchanger, but extracorporeal PFC gas exchange
is not an
S absolute requirement in many situations using MMLV.
If very rapid rates of heat exchange are desired, 10 to 20 mL/kg of PFC is
loaded into
lungs already filled to ~1/3 of total lung capacity (equal to ca. 30 mL/kg or
~FRC). This
brings total PFC load to a maximum of less than 2/3rds of total lung capacity
(2/3rds of 90
mL/kg = 60 mL/kg), and typically less than 60% of total lung capacity. As soon
as the
liquid has been loaded, it is then unloaded again, as rapidly as possible, to
FRC. Heat
transfer from liquid is rapid enough with MMLV that use of deliberate liquid
dwell times
are usually not efficient. Liquid loading and removal rates will be discussed
below, but for
maximal heat transfer rates, liquid alveolar "minute volumes'' (liquid
alveolar ventilation)
typically averages 25 to 33% of those for gas.
As liquid is loaded into the lungs, the volume of gas delivered with each
mechanical
breath is decreased, using a pre-determined peak inspiratory flow pressure as
the cut-off
point. Current experience with a canine model indicates that the maximum peak
airway
pressure which is tolerable without inducing significant volu- and baro-trauma
to the lungs
is ~-40 cm water (29 mmHg or Torr). These pressures occur at the ends of small
(10 mL/kg)
gas breaths, when the lungs already hold 40 to 50 mL/kg of PFC. The occurrence
of higher
positive airway peak pressures can be particularly avoided by use of small ( <
10 mL/kg)
PFC infusions at rapid rates (> 50 mL/kg/min), but these rapid techniques
require machine
control of infusion and suctioning (see examples below).
As liquid is suctioned from the lungs after the rapid phase of heat exchange
has
occurred, the tidal volume of gas delivered to the lungs should be increased
as the liquid is
progressively removed. When the lungs are fully loaded with PFC to 60 to 70%
of total
lung capacity, mechanical gas breaths must be very small in volume to prevent
overpressure. When the lungs are fully unloaded of easily suctioned PFC, gas
ventilation
can occur at a maximally effective rate and volume needed for C02 removal. In
MMLV,
particularly during cooling, C02 is removed by both PFC and gas ventilation.
PFC
infusion/suction cycle ratios may be adjusted, however, to increase C02
removal in
situations where this is especially needed. Increasing the time during which
the lung is
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unloaded of fluid allows for best time efficiency in C02 removal. at a given
maximal
airway pressure.
Heat transfer considerations suggest that most of a fresh load of PFC
delivered to
the lungs reaches the small airways after a few gas ventilations, since the
rate at which this
fraction of PFC reaches thermal equilibrium with previously infused PFC, and
with blood
flowing through the lungs, has been found to be very rapid, with the
equilibration half
times being much less than 1 minute. Again, this is probably largely due to
the assisted
connective mixing of cold and warm fluid due to gas-induced stirring seen in
MMLV.
Rapid-phase heat equilibration is complete by the end of PFC loading at rates
of loading up
to 6 mL/kg/min, since the average temperature of the expired PFC after a
loading cycle
does not increase, whether zero dwell time is used or an additional dwell time
of up to one
minute after loading is allowed. At these rates, average temperature of a
removed PFC load
remains constant at about 3°C below the animal's core temperature
during short dwell
times, because of the contribution of the dead space volume of cold liquid in
the large
1 S airways, which liquid requires a much longer equilibration time to warm
(see Example 2,
FIG.3 below).
At much faster PFC infusion rates, from 16 to 45 mL/kg/min (see FIGS and
FIG.9),
and loads of 23 mL/kg, the temperature difference between end-expiratory PFC
and animal
core temperature increases from approximately 5°C to 7.5°C. This
is presumably due to
incomplete temperature equilibration within body thermal compartment system
#1, which
is composed of blood, lung tissue and PFC in the smaller airways (see
explanation of
thermal compartments below). However, this effect is small within this range
of thermal
load, and these figures indicate surprisingly small decreases in heat transfer
efficiency
(from PFC to subject, or vice versa) with increasing total rate of PFC
"lavage." In this
series, using MMLV, efficiency varies from 92% at 6 mL/kg/min, to 77% at 50
mL/kg/min
(30 second infusions), to 57% at 50 mL/kg/min ( 10 second infusions).
A time of 30 seconds for infusion, and 15 to 20 seconds for liquid removal is
typical
for experiments at our fastest manually controlled liquid loading and
unloading rates (see
experiment 7 below). A typical algorithm for manually controlled MMLV for a
medium-
sized dog (20 kg) is to load PFC at 40 to 50 mL/kg/min and unload PFC at
approximately
80 mL/kg/min. For faster loading and unloading rates, allowing lower pressures
and more
time for gas ventilation, machine control of PFC infusion and suction is
required. FIG. l0a
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shows cooling rates for 3 animals cooled with machine infusions of 8 mL/kg (50
ml/kg for
seconds) followed by machine-controlled suction of the same PFC volume, for 4
to 7
seconds. These animals are compared in cooling rate with 5 animals cooled
using the
identical infusion rate (SO mL/kg/min), but with 30 second loads and
comparable suction
5 times. All of these animals received essentially the same amount of cold
fluorocarbon
during 18 minutes, but using infusions of FC of 1/3'd the size resulted in
about a 25% loss
of efficiency in heat transfer. Some of this loss occurred in non-insulated
liquid infusion
lines, where heat leaks between infusions became important for short
infusions, and induce
a kind of "heat dead-space," akin to the ventilator gas line dead space which
becomes
10 important during small tidal volume ventilation.
Theoretically Expected Cooling Rates
The PFC employed for most of the experiments reported here is 3M Company's
"FC-75," a proprietary mixture of fluorinated hydrocarbons reported to consist
mostly of
perfluoro-butyl-tetrahydrofuran (CF3-CF2-CF2-CF2-C4F80), and which has a
volumetric
specific heat capacity of 0.45 cal/mL/°C. A loading rate of 45
mL/kg/min (30 seconds)
followed by an unloading rate of 80-90 mL/kg/min gives a recycled liquid
minute volume
of 22.5 mL/kg infused and removed, averaged out over 45 seconds (time for
liquid entry
and removal). This results in an average PFC recycling rate of 22.5 mL/kg =
(.75 min) = 30
mL/kg/min (we will detail a number of exmples below with comparable infusion
and net
recycling rates but differing infusion loads). Since cold PFC enters the
animal at 3°C and
leaves at about 28-33°C (as the subject cools from 38°C to
33°C), this represents heat
removal at a rate of 30 mL/kg/min x 0.45 cal/mL/°C x 27°C = 365
cal/kg/min = 0.405
kcal/kg/min. For an animal with a mean body specific heat capacity of 0.70
kcal/kg/°C this
gives an expected cooling rate of (0.365 kcaUkg/min) = (0.70
kcal/kg/°C) = 0.52°C/minute.
In practice, maximal brain/tympanic cooling rates of up to 0.43 C/min have
been achieved
with this technique (Example 7, FIG 10), and faster rates are in theory
achievable with
colder PFC and faster PFC infusion. In our studies we typically cool animals
to ait
equilibrium temperature 5 to 10 C below initial body temperature (representing
an initial
drop in core temperature of 7 to 12 C). Rates of ~0.5°C/min represent
about I/2 of the
cooling rates available on cardiopulmonary bypass (CPB), but are superior to
cooling rates
achieved with all other techniques, both invasive and noninvasive. They are
also, as a
matter of practicality, far superior also to "time from intent to treat" CPB
rates, since times
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to achieve machine connection for the average patient must here be
realistically added into
the equation.
The fastest cooling rates reported here are 3 times the rate achieved by
Shaffer's
group in TLV cooling of cats (Shaffer TH et al.. 1984). Note the fact that the
fastest tested
rates of cold PFC recycled with MMLV. as detailed below (30 mL/kglmin
recycled) are
less than I/2 of those typically employed in TLV as applied by Shaffer et al.
at their coldest
liquid infusion temperatures of 10 C (minute volumes in Shaffer's study are
~56
mL/kg/min), and yet the cooling rate achieved with the MMLV technique is far
greater.
This is due to the relatively poor heat extraction and poor temperature
equilibration of PFC
in the lungs with TLV, especially at higher TLV rates, causing effective loss
of cooling
efficiency and power. This is suggested in Shaffer. et al. ( 1984) by the
large 12°C to 15°C
difference between their experimental animals' body temperatures and the
temperature of
their expired PFC. (Moreover, as previously noted, this difference should have
been even
greater, given the unexpectedly slow rate of body cooling reported).
At lower ventilation rates in TLV, cooling and heating efficicncy improve
somewhat, but now C02 removal becomes problematic. Heat removal never
approaches
maximum efficiency in TLV because the needed ventilation rate in TLV is
apparently too
high to allow good heat equilibration between the PFC already in the
functional residual
capacity of the lungs, and the cooled or warmed PFC delivered in a given tidal
volume
breath. The process of PFC heat equilibration and heat transfer in TLV is
convective and by
mechanical fluid transport (ie. circulation. bubble or vibration mixing) for
fluids on larger
scales (large airways) but increasingly by conduction at small-scales (small
airways and
alveoli). Convective heat transfer in small airways in TLV is impaired by
laminar flows
which inhibit the warmed or cooled PFC, which has been delivered into the
large airways,
from being mixed rapidly with the PFC already present in small airways. Mixing
(and
convective heat transfer at smaller scales) is greatly facilitated when micro-
bubbles of gas
(from a gas ventilatory breath) initiate small airway liquid turbulence. The
solution to the
problems of heat and gas exchange in liquid ventilation is therefore to
introduce gas into
liquid ventilation heat exchange systems (thereby creating micro-bubble small-
scale non-
thermal fluid mixing), and also to vary PFC and gas delivery independently to
the lung, so
that each can be used at their maximally efficient and least traumatic
delivery and removal
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rates. This is novel. A description of this general method is described below
with preferred
embodiments.
Description of Practical Protocol for MMLV
In MMLV, the liquid loading sequence is typically achieved by pumping the PFC
S from a reservoir through a circuit consisting of a heat exchanger, a 20
micron pre-filter
followed by 0.2 micron absolute filter, and an oxygenator (FIG.6A). For PFCs,
the
oxygenator, if used, must currently be a true membrane oxygenator, as the
current
generation of hollow fiber capillary oxygenators available in the U.S. are
fenestrated at the
micro-scale, and depend on the high H20/gas surface tension to keep gas and
liquid in the
oxygenator separated, and do not work with the relatively low surface-tension
PFCs. After
passage through the oxygenator, the PFC liquid is then delivered to the
subject through a
tube which is passed through the suction port down the endotracheal tube, to
the level of
the carina (FIG.6B). This single-lumen tube may be used to both deliver and
suction liquid
from the lungs. In the present implimentation, this tube is a specially
constructed flat-wire-
reinforced ultra-thin-wall tube with a fenestrated open-end, but many other
designs are
possible within the claimed type of device. Control of PFC infusion and
removal may be
done by a computer-controlled valve-driver and pressure-sensor device. Such a
device,
used for the series of illustrative examples herein, was designed by and built
by Korr
Biomedical Corporation, Salt Lake City, Utah, under direction of the patent
claimants.
In MMLV, mechanical ventilation with gas proceeds in routine fashion with a
pressure-limited or pressure-controlled mechanical gas volume-ventilator. This
includes
appropriate monitoring and control of peak inspiratory pressure, peak flow,
tidal volume,
minute volume, Fi02 (inspired oxygen concentration), gas composition, and
other relevant
ventilatory parameters. The mechanical ventilator in the best implementation
of the
technique must be able to sense and adjust these parameters appropriately as
liquid is
loaded and unloaded from the lungs. An additional relief valve must be
included in the
vacuum circuit also, in order to limit negative airway pressures which occur
after liquid
suction is completed during each cycle, and the suction circuit pump suddenly
becomes
exposed to airway gas.
The relationship between the volume of liquid in the lungs at any one time
during
mechanical gas ventilation, and the other variables of peak pressure, peak
flow, tidal
volume and ventilatory rate, is very important:
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1 ) As a first approximation, peak positive airway pressure should not exceed
40 cm of
water at the endotracheal tube external connector, and peak gas flow should
decrease from
a maximum of 60 L/min (LPM) to a minimum of 20 LPM or less, as the lungs are
progressively loaded with PFC. Similarly, tidal volume of delivered gas will
decrease to
essentially zero whenever the lungs are loaded with PFC to 50-60% of TLC (ca.
SO mL/kg)
and return to normal delivered tidal volumes of gas (i.e., 10 to 20
mL/kg/breath) when the
lungs are unloaded of fluid to a volume approximately that of functional
residual capacity
(30 mL/kg). For mechanically-driven systems, rapid (4/minute) small infusions
of fluid <
10 mL/kg may be used to keep maximal airway pressures low.
2) The tidal volume of gas breaths will decrease to zero or near zero when the
lungs are
fully loaded with liquid, in order to avoid baro-trauma and volu-trauma.
Generally, less
than half, and as little as one-quarter. of total lung fluid load may be
available for suction
removal, before suction pressures become unacceptable due to loss of fluid
return and
exposure of the large airway gas to high negative pressures. In a fully
flexible system, gas
ventilatory frequency may also decrease during times of high lung liquid
content, in order
to minimize gas flow pressures. During liquid unloading, and for a pre-
selected period
afterward, gas ventilation rate and volume will increase in frequency as
rapidly as the pre-
selected limiting peak inspiratory pressure (or peak flow pressure) will
allow, to a rate of 12
to 15 breaths per minute. This ventilatory rate has thus far been found to be
effective for
unloading of the lungs with PFC, via active suctioning. However, rates of up
to 30
breaths/minute and more have been used, and in some circumstances are
effective. High
frequency ventilation may also be used. In this implementation, a catheter
under the surface
of PFC liquid collecting in the bronchi may be involved in liquid removal, and
have little or
no effect on gas ventilation given around the removal catheter, until the
suction tip becomes
exposed to gas.
The reason for observed gas-ventilation facilitation of liquid removal seems
to be that
gas ventilations cause liquid in small airways to be moved to large airways,
where liquid
can be removed by active suctioning at the bottom of a gravity drain. In a
sense, MMLV
allows PFC to be both pushed and pulled out of the lung, without chest
compression.
Briefly, PFC can be rapidly suctioned from the lungs when ventilator gas
actively replaces
it.
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In TLV (prior art), the necessary ventilation frequency, which is in excess of
3 liquid
breaths per minute, does not allow the bubble-free PFC fluid. with its
limiting viscosity and
flow, to become available for removal in this way, and consequently in TLV,
tidal volumes
and "minute liquid volumes" (dV/dt) must be severely limited at higher
ventilatory rates.
By contrast, in MMLV the much higher gas ventilation rate at which fluid
removal rate
limitations occur allows for considerably better C02 removal than occurs in
total liquid
ventilation. This is a novel feature of the invention.
Consideration of MMLV Circuit Mechanics
Following delivery of a PFC load to the lungs, the liquid must be removed and
recycled through the heat exchanger/filter, and often oxygenator assembly.
This assembly
rids the liquid of C02, adds oxygen, removes mucus, bacteria and other airway
debris, and
appropriately heats or cools the fluid. Again, rapid but gentle removal of
liquid from the
lungs is best achieved by applying suction to the PFC delivery catheter (or
alternatively the
endotracheal tube) while imposing mechanical gas ventilation through a
different
concentric tube, per the algorithm described above.
External pressure (chest compressions) can also facilitate removal of PFC from
the
lungs, but under certain circumstances can also cause small tears or
lacerations in heavier,
liquid-loaded lungs. These tears are usually unimportant clinically unless
they result in
subsequent PFC in the pleural space ("perfluorothorax") in systems where the
PFC used has
a high vapor pressure (e.g., >40 mml~ig at normal body temperature, e.g., 3M
Company's
"FC-84" mixture). This so-called "perfluorothorax," when it occurs, is
clinically far less
problematic with low vapor pressure (e.g., ~10 mmHg at body temp) PFCs, such
as 3M
Company's "FC-40" mixture, or the proprietary pharmaceutical
perfluorooctylbromide, or
perflubron (LiquiventT"', Alliance Pharmaceuticals). Our invention of MMLV is
suitable for
use with all of these PFCs, and many others. However, for MMLV combined with
manual
or mechanical CPR, which greatly increases the risk of perfluorothorax, only
low vapor
pressures PFCs are suitable for use. Certain PFCs with undesirable low-
temperature
characteristics leg, perflubron) such as high viscosity or freezing behavior
near 0 C, cannot
be used as cooling media in the most extreme cooling applications described
herein.
However, low liquid temperatures (near 0 C) per se do not seem to cause
serious lung
damage, and animals have survived 30 minutes of rapid cycling of PFC cooled to
below 5
C, with core temperature drops of as much as -12 C. All animals allowed to
survive the
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most rapid cold PFC cycling temperatures and rates described herein, returned
to normal A-
a gradients (a measure of, lung oxygenation function) within 48 hours after
the procedure.
Appropriate ventilating gas composition is also very important for different
clinical
problems. In some situations, very high Fi02 will be desirable, while in other
situations
S some gas breaths may contain no oxygen for a short period. An example of the
latter would
be in cases of re-perfusion after cardiac arrest, where oxygen might be
temporarily withheld
for a few seconds in order to protect against exacerbation of re-perfusion
injury by
reintroduction of oxygen while free radical scavenging drugs are delivered to
the tissues.
The addition of drugs to the ventilating gas to overcome vasoconstriction of
the
alveolar and other lung blood vessels and/or to correct V/Q (V/Q = dV/dt /
dQ/dt)
mismatch, is also a feature of this invention. In particular, the addition of
nitric oxide (NO)
to the ventilating gas, or to the PFC via the oxygenator, in concentrations
ranging from S to
80 parts per million (or to effect) can overcome cold-induced vasoconstriction
of the lung
airways, and allow for improved rates of gas exchange.
1 S The following are representative data from canine experiments done by the
claimants and 21st Century Medicine (21CM), [Critical Care Research, IncJ
which
directly illustrate the performance as compared with other techniques which
exist in the
literature. Each experimental Example is a single exemplary animal, or a group
of
similarly treated animals, and will be used to illustrate one or more
principles of the
invention. They will also illustrate novel physiologic principles necessary to
understand
the mechanism of the invention. Although other materials and methods similar
or
equivalent to those described herein can be used in the practice or testing of
the present
invention, the preferred methods and materials are now described.
21" Century Medicine [now Critical Care Research] is an USDA licensed animal
2S research facility. All animals used in the work described herein received
humane care in
compliance with the Guide for the Care and Use of Laboratory Animals published
by the
National Institutes of Health (NIH publication 8S-23, revised 1996) and in
accordance
with all State and Federal laws regulating the use of animals in biomedical
research.
In deliberate departure from SI nomenclature, heats are often given in this
document in calories (1 cal = 4.184 J) or kilocalories (1 kcal = 4184 J). This
is done
because it greatly facilitates intuitive understanding of heat capacity
relations in liquids
and in watery solids, such as tissues.
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Example lillustrates the rapid rate of heat transfer which typically occurs in
Partial
Liquid Ventilation (PLV} when a load of warm or cold fluid is infused as well
as the
three compartment heat reservoir model.
EXAMPLE 1
Analysis of Partial Liquid Ventilation
Illustrated in FIG.1 is a canine model in which one load of non-isothermal PFC
is
given. In this experiment, one infusion of I.65 L of cold, (4.4°C)
oxygenated PFC (here,
the 3M Company PFC product "FC-84") is given rapidly (683 mL/min, over 145
seconds
= 31.3 mL/kg/min) by catheter passed below the lower tip of the endotracheal
tube of a
21.8 kg dog. The volume infused (75.7 mL/kg) is chosen to be near 80 % of the
total lung
capacity. Mechanical gas ventilation is continued through the entire
procedure, resulting
in very different thermal kinetics than in TLV (see discussion). For
simplicity of
interpretation, suctioning in this experiment is delayed until after a
suitable dwell time
(after complete temperature equilibrium had been reached).
FIG.1 shows a graph of canine temperatures in experimental Example 1. The
graph
shows temperatures over time, as 75.7 mL/kg of PFC at 4.4°C is loaded
into the lungs
over 145 seconds. Illustrated are animal esophageal temperature and indirect
brain
temperature as measured via a copper/Constantan thermocouple probe placed in
the
esophagus, and another on the right tympanic membrane. Also illustrated are
venous and
arterial blood temperature which were measured, respectively, by a thermistor
in a
thermodilution catheter inserted via the femoral vein into the inferior vena
cava, and by a
thermocouple probe inserted via femoral artery into the descending aorta.
Instrumentation of the animal as described is illustrated in FIG 1 a.
The FIG 1. temperature graph illustrates blood temperature (both venocaval and
aortic) falling and then rising, as PFC is infused, then allowed to remain in
the lungs.
Esophageal temperature, which is a marker for the temperature of the
relatively small
amount of PFC in the trachea and large airways, does not come into equilibrium
until
about 450 seconds after infusion of cold PFC ceases, with a half time of about
120
seconds. Aortic blood temperature reaches a nadir at essentially the time
infusion of cold
PFC stops, then rises with a much faster half time (about 80 seconds), with
equilibrium
complete at about 210 seconds after infusion ceases. Venous blood follows a
similar
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pattern with a smaller deviation, with the venous temperature nadir offset
from that of
arterial blood by 30 seconds - roughly the mean blood circulation time of the
animal.
The Three-Compartment Heat Reservoir Model
In these experiments involving anesthetized and paralyzed dogs which do not
shiver, the heat transfers are too rapid to show any effect from metabolism or
surface cooling, and thus only temperature effects from the chilled PFC
infused are seen.
In Example 1, about 2/3rds of the brain/tympanic temperature change of the
animal
occurs during the "dwell-time" of the liquid (time between completion of
liquid loading
and liquid unloading). This indicates that heat is transferred between lungs
and blood
very rapidly on this time scale (i.e., this system of PFC, lung parenchyma,
blood
volume, and probably some heart/arterial intima and muscle, comprise a single
thermal
reservoir or compartment for purposes of analysis, at this time scale).
Heat is then transferred more slowly (i.e., slowly enough to easily see on
this time
scale), and in an approximately exponential fashion, from this first thermal
compartment
(PFC, lungs, blood volume, etc.), to the next 2 compartments, as illustrated
in FIG 2.
Thermal equilibration with one of these compartments, comprising certain well-
perfused
tissues of the body, presumably the kidneys, heart, brain, liver and some
muscle, is
much more rapid, by a factor of about 6, than to the other compartment. These
tissues
therefore are taken to comprise one of two more thermal compartments
(Compartments
#2 and #3).
The fact that aortic blood temperature immediately (within lung-aorta
circulation
time of a few seconds) begins to rise after PFC infusion is stopped, indicates
that even at
these high loading rates of 31 mL/kg/min of PFC, there is no significant
difference in the
temperature of PFC residing in the small airways of the lungs, and the
temperature of
blood which has circulated through the lungs. (This fact is also indicated by
the relative
insensitivity of end-expiratory PFC temperatures to PFC loading and removal
rates, in
Examples to follow). A lag between the start of warming of aortic blood and
the halting
of cold PFC infusion, would indicate that a significant amount of infused PFC
remained
during this delay-time, which had not yet equilibrated with pulmonary
circulation.
Instead, at least up to the infusion rates of cold PFC tested here, a state of
continuous
near-thermal equilibrium between blood and most of the PFC in the lungs
(excluding that
PFC in large airways) is inferred. This is the rationale for including well-
stirred small-
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airway PFC, and blood, in the same thermal compartment in this model
"Compartment
#1 ").
The final equilibrium change in body temperature of 1.6°C suggests a
total animal
specific heat capacity of (75.7 mL/kg x 0.45 cal/mL/°C x 31.2°C)
= 1.6°C = 0.66
kcal/kg/°C. This is a reasonable value. Values for lean tissue and
blood are respectively
0.87 and 0.88 kcal/kg/°C (Cooper and Trezeh, 1971), and a lower mean
value for whole
animals reflects the lower water content and specific heats of bone and fat.
Compartment One to Compartment Two Heat Transfer
The nadir of aortic blood temperature (occurring at the end of the PFC load)
in this
experiment is S.5°C below body temperature, and the temperature for
venous blood at
this time is 2.6°C below initial body temperature. Tympanic temperature
at this time has
fallen 0.5°C. From these figures it is possible to make some simple
quantitative estimates
of heat transfer dynamics between heat compartments in this model, without the
use of
complex mathematics or computer modeling.
The 75.7 mL/kg of PFC given in this experiment represents a heat deficit of
(37°C
- 4.4°C) x 0.45 cal/mL/°C x 75.7 mL/kg = 1110 cal/kg. Of this
heat, at the time of the
end of the PFC load, the arterial blood (assumed on the basis of average land
mammalian
body composition to be about 20 mL/kg) holds approximately 0.88
cal/mL/°C x 20
mL/kg x 5.5°C = "100 cal/kg heat deficit, and the venous blood, 40
mI,/kg (by similar
calculation) holds about the same. The remaining PFC in the lungs, which can
be
assumed to be mostly at the same temperature as the arterial blood (see
argument above
and experimental data below) still has a heat deficit of 75.7 mL/kg x
0.45°C x 5.5°C = '
190 cal/kg. In '150 seconds, the PFC load has transferred (1100-190)/1100 = '
83% of
its heat deficit to Compartments #1 and 2.
At this time, body temperature has dropped by 0.5/ 1. 8 = 35.7 % of its final
amount, indicating that '1100 cal/kg x {1- 28%) _ '750 cal, is still to be
accounted for
at this point. As calculated, 390 of these "deficit calories" are in blood and
PFC. The
remaining -360 cal/kg heat deficit must therefore at this point be distributed
in internal
viscera very well equilibrated with blood, probably primarily in lung
parenchyma) tissue
and vascular tissue (including some of the endocardium). The mass of this
immediate
visceral heat sink, which is at arterial/PFC temperature of -5.5°C
below body
temperature, and assumed to have a typical specific heat capacity of organs
and lean
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tissue of 0.87 cal/kg/°C, can be estimated to have a mass of 360 cal/kg
= 5.5°C = 0.87
cal/g/°C = 75 g/kg, or 7.5 % of body mass. This non-blood "rapid heat
equilibrium"
visceral mass is thus about the same mass as the blood volume (also usually 7-
8% of body
mass), and in a 21.8 kg dog would amount to 1.6 kg of tissue. Thus, the "rapid
equilibrium compartment" or first thermal compartmen t of a dog receiving cold
PFC
lung lavage, is found to consist of about 50% of contribution from less than 2
kg of
certain internal viscera (probably mostly lung, with some heart and vessel
contribution),
and 50% from infused PFC, plus total blood volume. This compartment, since it
has
twice the heat capacity of the visceral contribution just calculated, would
comprise (2 x
0.075 x 0.87) / 0.66 = 20% of the heat capacity of the animal, and a bit less
than the
same fraction of the animal's mass.
Third Compartment Equilibration (Core to Peripheral Tissue Heat Transfer)
An initial 1.8°C hrain/tympanic drop in one-load cooling (implying mean
specific
heat capacity of only 0.59 kcal/kg/°C), is unreasonably high for a non-
obese, young
animal, and must reflect cooling of only part of the animal (i.e., the first
two thermal
compartments discussed above, plus only a fraction of Compartment #3) in such
experiments. Final equilibration of the first two thermal compartments with
the third
takes place after blood and tympanic temperatures equilibrate (Compartment # 1
and 2
equilibration) after about 4 minutes from the end of the PFC load (calculated
half time is
1.3 minutes). Tympanic cooling reaches a nadir of -1.8°C below the
starting body core
temperature at this time, reflecting Compartment #1 and #2 equilibration, then
slowly
rises toward a point -1.6°C below initial temperature, over an
additional 10 minutes
(calculated half time is " 6 minutes).
The delayed rise in body temperature after central cooling (which may be
thought
of as "after-rise" by analogy to "after-drop" during peripheral re-warming)
indicates
thermal equilibrium of compartments #1 and #2 is being reached with the third
thermal
compartment, which is not as well perfused as the brain and central body
organs. This
third compartment is assumed to represent fatty tissue and some cold-
vasoconstricted
muscle and skin, and is sometimes referred to in the literature as the
"peripheral tissues"
(vs. the better-perfused "core tissues. "). The third (or peripheral)
compartment may also
be expanded to include PFC in large airways (see the esophageal temperature
curve in
FIG. 1) which also warms with a similar, though slightly faster, half time.
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Estimation of the comparative size of the third thermal compartment (periphery
1
tissues) in the dog can be made from the "after-rise" in body temperature
which occurs
from the time the first two thermal compartments ("core tissues") reach
equilibrium
(about 7-10 minutes into the experiment), and the time the final "equilibrium"
body
temperature is reached at ' 16 minutes. During this time the increase in body
temperature
from 1.8°C to 1.6°C below the initial temperature indicates that
the heat capacity has
increased by a fraction 1. 8/ 1.6 =1.12 = i 2 % . This increase is due to
residual heat
capacity of the third compartment being made available, even though because of
its own
cooling half time, the fraction of its heat capacity left will by this time (7-
10 minutes) be
only about 1/3'° of its original, since this time represents about 1.5
half times for this
process, meaning the 3'd compartment is about 65 % of the way to equilibrium
with the
lungs by the time it begins serving as the sole warm reservoir for rest of the
body.
The 12% increase therefore represents about 12/.{1-.65) = 34% of the initial
heat
capacity of the animal, as a crude and approximate estimate. This leaves the
remaining
46% for compartment #2 (see FIG .2). Note that '/4 of the 20% heat capacity in
Compartment #1 is PFC, and for smaller loads on top of FRC PFC content, this
number
will tend to be closer to 17
See also FIGS. 8, 9 and 10 for examples of this phenomenon of "after-rise." In
summary, we find that this animal can be modeled as three thermal
compartments, with
heat or cold transferred from the first to the second rapidly (with a half
time of about 1.3
minutes) and then from both of these compartments to the third compartment at
a slower
rate. Half time for this process is '6 minutes in this illustration and also
in Example 6',
FIG. 9. It is '6.5 minutes at the end of PFC infusion in Example 3, FIG. 5,
illustrating
equilibration with a colder animal. Note that in the literature, equilibration
between core
and peripheral compartments is quite dependent on absolute temperature, since
it is
greatly influenced by cold induced vasoconstriction and drop in total cardiac
output.
Because of low blood flow with severe vasoconstriction, movement of heat
between core
and periphery in deep hypothermia is slow enough to be mediated partly by
direct thermal
conduction through solid tissues, in addition to blood circulatory heat
transport.
Estimate of the Quantitative Rate of Heat Transfer
In the above experiment (Example 1), the rate of heat transfer from
Compartment
#1 to Compartment #2 and #3 can be estimated by noting that the maximum rate
of
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temperature increase in aortic (arterial) blood temperature is 1 °C in
0.5 minutes (2
C/min), suggesting a maximum heat removal rate for the arterial blood lung
PFC, and
most of Compartment # 1, which are presumably at about this temperature also.
If 15
of the heat capacity of the animal is in non-venous blood in compartment #1,
then rate of
heat transfer represented by 2 C/min is 2°C/min x 0.15 % x 21.8 kg =
(0.64 kcal/kg) _
5.1 kcal/min = 356 watts. If the heat transfer by blood at this time is
crudely estimated
by the Fick principle (using thermodilution) from the estimated cardiac output
(about 2.5
L/min for an animal this size) and temperature difference between venous and
arterial
blood (3°C), the maximal blood heat transfer power at this time is: 2.5
L/min x 0.88
kcal/°C/L x 3.0°C = 6.6 kcal/min = 460 watts. These numbers
suggest that blood heat
transport alone is more than enough to account for heat transfer from
Compartment #1 to
Compartments #2 and #3 in this Example.
Indeed, in later Examples (see FIG ) we find that the fastest PFC rates
employed
(infusion plus suction (30 mL/k/min) typically give PFC delta T's of 17 C, for
calculated
cooling powers of about 4.8 kcal/min = 336 watts. These typically cool 20 kg
dogs at
an average net rate of 0.33 C/min, implying (assuming heat capacities of 0.7
kcal/kg),
cooling powers of 20 x 0.7 x .33./ 88 = 4.6 kcal/min = 322 watts.
Arterial/venous
temperature differences of 1 C going to 1. S degrees over the cooling run are
common
(see FIG ) and imply cardiac outputs by thermodilution of 4 liters per minute,
dropping
to 3 L/min Again, blood transport of heat seems entirely adequate to explain
most
features of heat transfer from lungs to body in this model.
Conclusions
There are several conclusions to be drawn from this example (Example 1), which
are important:
1) If gas breaths are used, PFC reaches very rapid thermal equilibrium with
blood, even
at PFC infusion rates greater than 30 mL/kg/min.
2) If gas breaths are used, heat or "cold" from an infused PFC load is rapidly
transferred
to a larger mass of lung, vascular tissue, and blood, which acts as a short
term storage
"heat capacitor" which has an effective heat capacity 3-4 times that of a 60
mL/kg (2/3rds
total lung capacity) load of PFC. This, combined with some heat transfer from
blood to
the peripheral tissues during even the fastest PFC lung loading, guarantees
that at least
75 % of heat or "cold" will be extracted from a significant fraction of a lung
load of PFC
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at infusion rates of '30 mL/kg/min, even with no deliberate "dwell time." For
example,
as seen in FIG. 1, this equilibrium is fast enough that when loading (to '80%
total lung
capacity volume) with PFC is accomplished in 2.4 minutes, less than 1/Sth of
the total
heat absorbing capacity of the PFC remains at the end of infusion.
S 3) Equilibrium cooling or warming of PFC in large airways (effective thermal
"dead
space") proceeds slowly, and the amount of heat-sink in this fluid is of
interest, because it
is lost during rapid ventilation. However, since it would take very slow
ventilation (or
infusion/suction) to equilibrate this fluid, an efficient protocol ignores it,
and uses no
dwell time. With gas ventilation augmenting mixing, the thermal load in PFC in
smaller
airways is very rapidly removed, and this (faster) transfer is too rapid to be
limiting to
heat transfer in the present model.
The slow esophageal re-warming in the previous experiment illustrates the
relatively slower thermal equilibrium of chilled PFC in large airways, even
though the
gas for mixing and warming PFC in smaller airways, is blown through the cold
PFC in
larger airways. This PFC remains cold because it is only slightly affected by
the small
heat capacity of gas blown through it, and because it is physically too far
from blood and
other tissues to be warmed at the same rate as PFC in the lung periphery.
Experimental Example 2 (FIGs 3, 4) illustrates heat transfer in a model with
multiple cycles of PFC loading and unloading. In each cycle, cold PFC is
loaded from
FRC (functional residual capacity) to about twice this volume, accompanied by
constant
gas rate ventilation. This protocol does not represent Total Liquid
Ventilation (TLV),
because ventilation and C02 removal are accomplished mostly by gas breaths
from a
conventional gas ventilator (here at a constant rate of 12 breaths/min).
However, gross
heat transfer in this experiment is accomplished by unloading the PFC liquid
soon (or
immediately) after loading, and then repeating the procedure with multiple
cycles of
freshly chilled liquid. This is similar to a multiple lavage procedure with
ice water in the
stomach or peritoneum, but has not been described as a technique for PFC
infused with
air into the lungs or other body compartments, and is novel when used as such.
Moreover, the use of gas ventilation mixing specifically and deliberately to
assist with
transfer of heat from PFC liquids to the lungs/blood, is novel.
The principles above may be used to design a maximally efficient protocol for
cooling a subject with MMLV. Example 2 illustrates one possible protocol.
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EXAMPLE 2
Illustration of Mixed-Mode Liquid Ventilation (MMLV)
Illustration of the concept:
In the following illustration (Example 2), a 25.7 kg anesthetized, intubated,
and
paralyzed dog was given chilled PFC (temperature -1°C) rapidly into the
endotracheal
tube, followed by removal by vacuum suctioning. FIG. 3 shows tympanic, rectal,
venous,
and aortic blood temperatures as a function of time. Also illustrated on the
same scale are
temperatures of the removed PFC with each cycle, measured in the suction
reservoir
(which is emptied after each cycle).
This experiment illustrates use of multiple loads of chilled fluorocarbon
liquid from
FRC to 2 x FRC, with varying dwell times after each liquid PFC load (here
again the
PFC is 3M Company's product "FC-75 "). In this illustration, the rate of PFC
liquid
loading is 155 mL/min, or 6.0 mL/kg/min, to a total of 24 mL/kg for each 4
minute load,
followed by suctioning over 2 minutes to remove fluid at a rate averaging
approximately
12 mL/kg/min. The relatively smooth cooling temperature curves are for the
entire
animal (tympanic and rectal temperatures), whereas the larger, periodic
temperature
fluctuations (total of 15) are blood temperature swings (as measured with
thermocouples
in the aorta and inferior vena cava). Each temperature fluctuation represents
a cold PFC
load, and subsequent removal by endotracheal suctioning.
At 6 mL/kg/min average infusion rate, rather than 31.3 mL/kg/min, initial
loading
in this experiment was done much more slowly than in Experimental Example 1.
After
the first two loads to approximately 60 mL/kg (approximately 2/3 FRC), an
additional
dwell time (in which PFC is neither infused or removed) of 6 and 4 minutes
respectively
was allowed for blood temperature to equilibrate with tympanic temperature.
During
these times, blood temperatures changed .by 4 to 5 times the rate of tympanic
temperatures, indicating that by this time (17 minutes into cooling) the heat
capacity of
the compartment in rapid thermal equilibrium with the blood (i.e., lung tissue
and
peripheral PFC in the lungs) was again relatively small (on the order of 20 %
), compared
with the heat capacity of the well-perfused, or core thermal compartment. At
this time
(17 min, at equilibrium after 2nd loading cycle), tympanic temperatures had
dropped by
only 1.5°C (0.09°C/min, or approximately 5°C/hr).
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In the following 13 cycles (15 cycles total), loading and unloading with
infusions of 24
mL/kg was continued (both at previous rates and volumes, thus making 4 minute
infusions), but dwell time was cut to 60 seconds, then 30 seconds, and finally
to zero for
the last 4 cycles. The effect of this change in dwell time on care cooling is
to smooth out
changes in temperature, but not markedly change the temperature of removed PFC
or
change the rate of cooling, as shown in FIG.3.
FIG.3 shows that average temperature of extracted PFC is about 3°C
below venous
temperature (venous temperature is assumed to be Compartment ~#1 temperature}.
Thus,
heat exchange within compartment 1 is seen to be substantially complete. Heat
exchange
is not complete between compartment 1 and 2 (venous blood and soft tissues)
when liquid
begins to be removed in this model, and this is a typical feature of heat
removal with
more rapid rates in MMLV.
Heat Transfer Measurement And Calculation of Heat Capacity
The amount of heat removed from the animal by the PFC may be easily assessed
by measuring the temperature of the PFC infused, and the temperature of the
mixed load
of removed PFC. In this example, mixing (30 seconds of mechanical stirring)
was
performed on the volume of suctioned PFC to alleviate temperature gradients in
PFC as it
is removed. Mixing allows for mass-averaged temperature and heat content to be
assessed.
FIG.4 (table of suction volumes and heat contents) shows temperatures of PFC
inspired and removed in this example, their temperatures, and the heat removed
with each
load of PFC infused and removed. Because some PFC remains in the animal at the
end of
the experiment, and because suction volumes do not always match infusion
volumes,
heats can be calculated in two ways. In the first method, heat infused and
removed is
calculated for all volumes of PFC, and then a correction is added for the
volume
difference between infused and removed PFC. This residual PFC is assumed to
remain in
the animal and to equilibrate with the animal's final body temperature. When
this is done,
total heat removed is found to be 125.9 kcal + 10.9 kcal = 136.8 kcal.
Alternately, heat
can be estimated by the difference in mean infusion and suction line
temperatures during
the middle of infusion and suction, and using infusion volumes only (on the
assumption
that suction volumes will be similar over the long run). This method, which
does not rely
on measuring suction volumes, is the only one which lends itself to heat
measuring
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during rapid unloading experiments, in which suction volumes cannot be easily
measured,
but in which infusion volumes are easily estimated from infusion rates. Such a
calculation
for Example 2 yields an estimated heat removal of 130.8 kcal, in good
agreement with
the more difficult method. Specific heat capacity for this animal is then
estimated to be
136.8 kcal/25.7 kg/7.5°C = 0.71 kcal/kg/°C. This is a reasonable
number, and suggests
that the method has no serious unaccounted-for heat losses at these low
infusion rates.
Example 3, shows the results when infusion and suction rates are increased.
EXAMPLE 3
Effect of More Rapid PFC Cycling
This experiment is similar to the one previously described, but with an animal
weight of 24.5 kg and infusion rates up to 410 mL/min = 16.7 mL/kglmin, and
increased suction rates.
In this experiment, the first two cycles were each loaded with 25 mL/kg PFC
over
4 minutes (6 mL/kg/min), then allowed a dwell time of 1.5 minutes. To
assessthe effect
of shorter cycle times and more rapid PFC turnaround, no dwell times were used
in
subsequent cycles, and infusions were shortened in cycles 3, 4, and 5, to 2.5
minutes (34
mL/kg load at a rate of 13.6 mL/kg/min) with 1 minute suction times. Finally,
infusions
were shortened to 2-minute loads at the fastest infusion rate of 16 mL/kg/min,
for loads
of 33.4 mL/kg. The effect on heat transfer is obvious from the increased rate
in body
temperature fall during cycles with faster infusion and zero dwell time.
Rectal
temperature fell a rate of 1.5°C/14 min = 0.107°C/min =
6.4°C/hr over the first two
cycles. Subsequently, more than doubling infusion rate and eliminating dwell
time
increased effective infusion + removal rate to 16.7 mL/kg/min x 2/3 = 11.1
mL/kg/min,
and an increased rectal temperature drop rate to 0.15°C/min =
9.0°C/hr.
At this liquid infusion rate (maximum 16 mL/kg/min - average of 11.1
mL/kg/min infused and removed), there is not enough time for blood
temperatures to
reach equilibrium with tympanic (brain) temperatures before liquid begins to
be
withdrawn, since half time for this process is again 40 to 60 seconds. At the
end of the
experiment, this is seen in a steady divergence of venous blood and solid
tissue
temperatures. However, this heat transfer half time provides more than enough
time to
transfer most of the heat out of the fraction of PFC in each infusion which
reaches the
lung periphery, as seen by the continuing small temperature gap between
expired PFC
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and venous blood temperature. This gap is thought to represent mostly PFC in
large
airways, as discussed above.
Calculation of total heat removed in this experiment was made by summing the
differences in suction and infusion temperature for each cycle, and
multiplying by PFC
S specific heat capacity and the infused volume. The total heat removed was
218.2 kcal or
8.9 kcal/kg. This, with a total temperature drop of 11.7°C, implied a
specific heat
capacity of 8.9/11.7 = 0.76 kcal/kg/°C. Again, heat is accounted for.
In this example, total PFC volume infused and recycled was 16,469 mL over 85
min, for an average cycle rate of 7.9 mL/kg/min. In order to obtain faster
temperature
descents of 20°C/hr, the goal in this series, loading and unloading
rates of 2.5 times this
rate would in theory be required. Unloading at these rates would need to be
facilitated by
ventilations during unloading, according to a gas ventilation algorithm, such
as that
previously discussed.
Accordingly, the next experiment to be illustrated involves full Mixed-Mode
Liquid
Ventilation, using a gas and liquid infusion algorithm.
Example 4 is a similar experiment using a 17.3 kg dog and different infusion
rates.
EXAMPLE 4
Full Mixed-Mode Liquid Ventilation with very rapid PFC loading and variation
in
ventilatory rate during unloading.
. This experiment (FIG.6) using a 17.3 kg dog illustrates the result when
infusion
rates are increased by a factor of 2.6, from 16.7 to 44 mL/kg/min. At these
rates, PFC
liquid loading to 1.5 x FRC is accomplished in approximately 60 seconds (to 44
mL/kg)
for the first load, and 30-40 seconds per load, once an equilibrium
distribution of PFC in
the lungs has been reached (beginning at cycle 4). This results in average
infusion
volumes of about 30 mL/kg, on top of FRC. Infusions were cut off when
pressures
reached 40 cm H20 = 29 mmHg.
These natural infusion and removal volumes at equilibrium cycling (30 mL/kg)
are
typical in MMLV experiments, though smaller infusion volumes ( < 20 mL/kg)
have
recently been found to provide much better control of peak large airway
pressures. In
order to remove liquid at rates comparable to infusion, it is necessary to use
a high flow
femoral cardiac bypass venous return canuia (such as the 19 French diameter
flat-wire
Biomedicus venous return canula used in this experiment) to remove PFC with
vacuum
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suction. In addition, it was necessary to control ventilation rates and
pressures
continuously during loading and unloading, both in order to avoid exceeding
peak
permitted airway pressures (40 cm of H20) at the end of loading cycles, and
also in order
to facilitate liquid removal during unloading, as described in Section I. The
result, in
which gas ventilation and liquid ventilation proceed independently,
constitutes the novel
invention claimed (Mixed-Mode Liquid Ventilation, MMLV).
With MMLV, PFC removal rates in Example 4 averaged 1390 mL/min early in the
experiment, and fell to 900 mL/min by the end of the experiment for the same
volumes
suctioned. These effects were attributed to increasing viscosity of the fluid,
and/or
differences in airway dynamics making less fluid available at lower body
temperatures.
The average time in which a 578 mL load of PFC could be both loaded and
removed in
this experiment, was 70 seconds at the beginning of the experiment, and 80
seconds at the
end of the experiment. This resulted in an average PFC cycling rate of 578
mL/75 sec =
7.7 mL/sec or 462 mL/min. For this 17.3 kg animal, maximum cycle PFC rate was
thus
462 mL/min ~ 17.3 kg = 26.7 mL/kg/min. This is 2.4 times the maximum infusion
cycle
rate (11.1 mL/kg/min) of Example 1. However, this did not reach the expected
cooling
rate of 21.6°C/hr (2.4 x 9°C/hr), primarily because the PFC was
infused at a higher
temperature in this experiment (3°C to 4°C). The actual cooling
rate in this experiment,
as seen in FIG.7, was 17.6°C/hr (rectal) and 22.6°C/hr right
tympanic (tympanic
temperatures are reflective of brain cooling). In this example, rectal
temperature was
decreased 9°C in 31 minutes. Tympanic and rectal temperature descent
rates differed
slightly due to non-equilibrium between the 2nd and 3rd heat compartments in
the animal,
as discussed above.
Estimation of heat removed in Example 4 was performed as in Example 5. Total
heat removed was estimated to be 123.4 kcal, with a body temperature change of
9.5°C.
Specific heat capacity was then 123.4 kcal/ 17.3 kg/ 9.5°C = 0.75
kcal/kg/°C.
Example
EXAMPLE 5
Scale-up of MMLV
In a larger animal (30 kg) subjected to MMLV using the same technique, but
scaling up liquid infusion flows, approximately the same cooling rate
(20°C/min) was
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achieved by MMLV (FIG.B). For this scale-up, liquid infusion was increased
slightly
more than linearly with animal mass (to 46 mL/kg/min), to make up for non-
linear effects
in PFC removal. However, at this rate we found that pressure-limited PFC
volume loads
were slightly less than in the smaller animal (Example 2) and averaged 740 mL
(24
mL/kg) in the experiment, with good regulation of maximum peak pressure. It
was found
later in this experiment that when pressure constraints were relaxed and peak
pressures
were allowed to increase to 50-60 cm H20, 1360 mLs could be loaded (45 mL/kg).
Suction rates in this experiment averaged only 840 to 900 mL/min (data not
shown),
mainly due to a poorer algorithm for managing ventilation. These results
illustrate the
sensitivity of MMLV to ventilatory algorithms and airway pressures (this
animal suffered
lung baro-trauma, but survived). PFC cycle time rate for this animal was only
11 mL/sec
or 663 mL/min. For a 30 kg animal this calculates to 22 mL/kg/min. The cooling
rate
actually achieved of 20.3°C/hr tympanic, and 18.5°C/hr rectal,
was similar to Example 4
(17.4 kg animal), despite only 83% of the PFC recycle rate/kg. This was partly
due to a
smaller lean body mass fraction in the 30 kg animal, with a reduction in
specific heat
capacity. Specific heat capacity for this animal was calculated, as above, at
0.70 kcal/kg/°
C.
In MMLV, infusion rates generally will need to be increased at a more than
linear
rate per animal mass. This is because average PFC fluid removal rates do not
scale well
with animal size, due to a maximum removal rate set by the removal catheter
and suction
system. Larger animals are expected to show slightly faster gross PFC removal
rates due
to longer times in which pure PFC, rather than a mixture of PFC and gas, is
available for
removal during suctioning. However, these effects may be partly offset by
better system
design. For example, in the 17.3 kg animal infused at 45 mL/kg/sec, removal
times were
SS % to 78 % of infusion time. For a 30 kg animal with comparable infusion
rates and the
same system, total liquid removal time had increased to 145 % of infusion
time. However,
in Example 6, FIG.9, a 19.8 kg animal infused at 46 mL/kg/min, improvements in
the
suction system allowed removal rates of 80 mL/kg/min (1600 mL/min), and
suction
times had decreased again to less than 70 % of infusion times.
Similar increases in the ratio of infusion to removal are expected for even
larger
animals, and in humans. Some of these effects are expected to be ameliorated
by the use
of larger diameter suction canulae in humans, and by improvements in the
liquid circuit
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which minimize flow resistance. For a 70 kg human, even if no improvements
over the
presently demonstrated system are realized, an infusion rate of 50 mL/kg/min
(easily
achievable), and a removal rate of 1600 mL/min = 23 mL/kg/min, will result in
a PFC
recycle rate of 50 mL/3.2 min = 16 mL/kg/min, which should allow for cooling
rates of
more than 50% of the demonstrated 24°C/hr at 30 mL/kg/min (i.e.,
approximately 12°
C/hr), depending on specific heat capacity of the subject. Even such lower
rates are
higher than previously reported for animals or humans using non-bypass cooling
or any
other method of external cooling likely attainable in the field. Moreover,
they are still
fast enough for neurological emergencies.
In Example 6, infusion volumes were cut down.
EXAMPLE 6
MMLV with increased suction rates
In Example 6, FIG.9, PFC infusion rate was maintained at 46 mL/kg/min (910
mL/min in a 19.8 kg animal), but infusion volumes were cut to 20 mL/kg (with
an
infusion time of 25 seconds). Suction rates were increased by engineering
modifications
which decreased resistance in the suction path. Effective PFC turnover in this
model
reached 27 mL/kg/min, and rates of cooling for the "core tissues" of the
animals
(Compartment #1 plus Compartment #2) reached 30°C/hr. This is the brain
cooling rate.
The animal was cooled only 5°C before infusions were stopped, and the
large 1.0°C
"after-rise" which then occurred (half time 6 minutes) illustrates the
difficulties of cooling
an animal very quickly with any method which uses the circulation as the chief
heat
transfer medium, such as MMLV or cardiopulmonary bypass. Peripheral
vasoconstriction
in very rapid central cooling always makes "overcooling" necessary, in order
to end at
the desired core temperature after final equilibration. In this example, a
cooling of 6.0°C
core/brain temperature would have taken about 12 minutes. This is the time
required to
remove enough heat to permanently cool the animal by 5°C. Uniform
cooling of 5°C over
12 minutes (1/5 hr) gives an effective cooling rate of 25°C/hr (5 C =
I/5 hr) over this
range. Specific heat capacity calculated in this experiment by previously
detailed methods
gives a value of 0.79 kcal/kg/°C.
Efficiency
In FIGs 3-6 it can be seen by examining venous and arterial temperature curves
that at
slower infusion rates the venous/arterial temperatre difference has a chance
to go to zero,
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or nearly zero, before new PFC is infused, indicating an unused or inefficient
length of
time in which heat is still being transferred from blood to body, but no
longer from lungs
to blood. This happens even at the fastest inspiratory infusion rates of 50
mL/kg/min, if
infusion cycles are too long due to overlong suction, or deliberate induction
of dwell time
(see temperature FIGS. 7 vs. 9 and 9a from Examples 4 vs. 6). Only with very
short
cycle algorithms does significant A-V temperature difference remain thoughout
the cycle,
although the time-integrated area of A-V temperature difference at this
infusion rate
reaches a maximum somewhere between a total infusion cycle time between 15 and
40
seconds. Even shorter infusion times begin to demonstrate a loss of efficacy,
because if
"residence" times for PFC in the lung are cut to the order of a few seconds,
it is possible
that any given infused volume of PFC can be removed from the lung before it
has gone
far enough toward reaching thermal equilibrium.
This maximum where area under the A-V temperature curve is largest represents
the
greatest heat transfer, and the best overall efficiency for this particular
infusion rate.
For all but the shortest loading times ( < approximately 15 sec), the most
successful
algorithm (in terms of rate of heat transfer) is generally to unload gas-mixed
PFC at the
maximum possible rate, directly after an infusion ends. Typical infusion
volumes load
PFC to 1/3 to 2/3rds total lung capacity. In some examples to follow, the
efficiency
problems of very short cycles will be illustrated.
EXAMPLE 7
Reliability of Cooling Rates In MMLV
FIG. 10 . shows body temperatures of a group of 5 animals (mean wt 20.6 kg)
subjected to MMLV with chilled PFC, under reasonably identical conditions.
These
conditions were chosen on the basis previous experiments to result in a rapid
core body
temperature drop of about 7 C, resulting in a final equilibrium temperature
drop of 5 C.
This was chosen to be indicative of an emergency hyperthermia situation, in
which
emergent cooling by more than 5 C would not be expected to be necessary. In
order to
achieve cooling rates close to the maximally obtained previous rates, an
infusion volume
was set at 50 mL/kg/min for all animals for 20 seconds (16.6 mL/kg), with
suctioning
controlled by pressure, and averaging about 15-I7 seconds (average cycle
length was 37
seconds).
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Animals began at normal body temperature, and were cooled with rapid cycling
of
PFC for exactly 18 minutes, with results shown. Average amount of PFC cycled
was a
total of 11.3 L per animal, or 550 mL/ kg. Body core temperature dropped by
7.3 +/-
0.4 (S.D.) C over 20 minutes, with final equilibrium temperature reached,
after thermal
compartment equilibration, of -6.0 +/- 0.13 (S.D.) C. The animals lost 4.2
Kcal/kg,
and the average temperature difference of fluid infused to fluid removed was
17 C., for
an average heat transfer efficiency of 60 +/- 8% . This temperature drop for
this neat
gave calculated heat capacities for the group of 0.73 +/- 0.1 kcal/kg/C.
In FIG. 10 the curve of typanic temperatures over time for this group of
animals is
compared with anaesthetized controls cooled by packing them in ice. Cooling
rates with
MMLV were not only 4 to 8 times as fast, but were also uniform and
predictable.
EXAMPLE 8
COMPUTER CONTROLLED MMLV
FIG. 6 is an illustration of a device for administration of Mixed-Mode Liquid
Ventilation (MMLV). Such a device, designed in prototype by Korr Biomedical,
Inc,
Salt Lake City, Utah, was built under direction of the authors, and used in
the following
illustrated examples. In this device, PFC is pumped and suctioned by the usual
methods,
but a computer actuates a set of valves which divert PFC from recirculation
from a
holding reservoir through a heat exchanger, into the animal instead. During
suctioning of
PFC, re-circulation of PFC continues through a bypass loop, so that the pump
may run
continuously.
The computer-controlled valve assembly allows infusions of cooled or warmed
PFC
into a catheter which is positioned in an endotracheal tube connected to a gas
ventilator.
PFC is also conditioned by being passed through a silicone membrane oxygenator
and
filters to remove respiratory tract secretions and bacteria. The computer
monitors airway
infusion, and has a cut off for high pressures on infusion and suction cycles.
In the best
implementation of the present invention, the ventilator is also controlled by
the same
computer, which monitors airway pressures and adjusts both PFC infusion and
gas
ventilation rate and tidal volume to insure that airway pressures do not
exceed critical
values. Pressure during suctioning is also monitored to insure that suctioning
halts when
liquid return slows, an event which is marked by peak inspiratory pressure
reaching 2-5
Torr. In the most general implementation of the device, PFC infusion can be
continuous,
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or it may be timed and pulsed with each inspiration. In the following
examples, PFC is
infused continuously.
FIG. l0a shows the result of using this method with 3 animals, as compared
with the 5
in the FIG. 10 discussed above. Although the machine-cooled animals received
very
nearly the same amount of cold PFC (650 mL/kg), and even more heat extraction
as
calculated from maximum inspiratory and expiratory fluid temps (5.3 kcal/kg),
they did
not cool as fast, as seen in the FIG. 10a. (net cooling of only 4.5 C). = 75 %
that of the
hand controlled 20 sec infused group in FIG 10. Unrealistic heat capacities
calculated
from these numbers (1.2 kcal/kg) makes it obvious that the problem is
temperature data
from the inpiration and expiration fluid which could not be fairly
approximated by a
rectangular integral. Assuming these animals had much the same heat capacity
as the
others, suggests that they were cooled only 75 % as efficiently, for a total
efficiency of
75 % x 60 % = 45 % . Evidently the average cycle time in these experiments (
15 second
- 10 seconds infusion, 5 seconds suction) was too short for the volume
infused. These
experiments represent the limits of infusion frequency for these infusion
rates.
EXAMPLE 8
Blood Gas Measurements May be Normal During MMLV, Even Without Cooling,
and At the Fastest Liquid Infusion Rates
In the experiments discussed herein, a variety of physiological parameters
were
monitored to establish the effects of MMLV on hemodynamics, gas exchange, and
lung
function both before (baseline/control conditions), during, and after the
procedure. As
previously noted, TLV, except under basal conditions, results in progressive
hypercapnia
with associated respiratory acidosis. Similarly, TLV is constrained as to the
V02 that can
be delivered to the animal in hypermetabolic states. The purpose of the
sophisticated
physiological monitoring discussed below is to assess the safety and efficacy
of MMLV
in inducing hypothermia. In addition to evaluating the animals' condition via
discrete
physiological measurements, the animals were allowed to survive for periods
ranging
from 24 hours to 2 weeks or more following the experiments in order to
evaluate the
effects of the procedure by clinical, laboratory, and histopathological
criteria.
Animals subjected to MMLV were instrumented for mean arterial pressure (MAP),
central venous pressure (CVP), mixed arterial oxygen saturation (Sp02) and
central
venous oxygen saturation (SV02). Arterial pressure was monitored via a line
inserted
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into the abdominal aorta via open cut-down of the femoral artery. CVP and SV02
were
monitored via a fiber optic pulmonary artery (PA) catheter advanced via open
cut-down
of the femoral vein to the level of the right atrium.
Arterial and venous blood samples were drawn before, during, and after the
procedure to assess the following parameters: pH, pC02, p02, sodium,
potassium,
chloride, ionized calcium, lactate, hemoglobin and hematocrit. For a number of
experiments (FIGs 11 and lla) central venous and arterial blood gas samples
were drawn
every two minutes during the active cooling phase.
The following respiratory parameters {gas phase) were monitored and acquired
in
some animals using a Novametrix Medical System (Wallimgford, CT) C02SM0
Respiratory Profile Monitor inserted into the ventilator circuit above the
endotracheal
tube. These parameters were: respiratory rate (RR), end-tidal pC02 (EtC02),
minute
C02 production (VC02), mean airway pressure (MAPa), minute volume (MV),
inspired
tidal volume (VTi), expiratory volume (VTe), peak inspiratory pressure (PIP),
positive
end expiratory pressure (PEEP), and peak inspiratory flow rate (PIF).
Respiratory Mechanics and Gas Exchange Data from Example 6
FIG. lla and llb show representative pa02 data from the experiments in FIG 10
and
10a, throughout a decrease in core body temperature of ~5°C, or when
being given body
temperature PFC. The pa02 at 100% 02 does not undergo significant fluctuations
when
cold PFC is given, but on room air, the A-a gradient typically falls by 20
Torr for a day
after PFC has been used, then recovers over the next day. For warm PFC
cycling, it is
interesting that even with oxygenated PFC, pa02 dropped signifcantly while the
dogs
breathed normothermic PFC, indicating a signftcant shunt and possibly V/Q
mismatch.
This, interestingly, was temperature dependent (the only variable). It was
seen in
automated and hand controlled trials. It remains unexplained, but perhaps
represents a
cold-prevented problem with vascular opening, which was not seen with body-
temperature PFC given at the same rates and pressures.
These data indicate that the volume of oxygen delivered to the tissues of the
animal is
adequate throughout the entire period of MMLV.
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Control of C02 Removal May Be Maintained Even at the Fastest Warm PFC
Infusion Rates
FIGs. llc and lld show pC02 direct blood gas measurements (NovaStat II
machine)
during the active phase of cooling in the same groups. Here the same trend is
evident,
S with cooled animals doing much better on gas exchange. Measurements of
directly
obtained arterial paC02 (femoral artery blood draws) show paC02 to be well-
maintained in
cold PFC cycling in MMLV, but that many of these protocols were not sufficient
to remove
C02 production when identical infusions of body-temperature PFC (38 C) was
used as a
control for the effects of chilled PFC.. Again, animals had no difficulty
building up C02
in either manually controlled or machine controlled MMLV cooling, but a large
difference was seen between dogs ventlated with body temperature liquid (no
more than
0.5 C higher than their blood temperature-not high enough to increase
metabolic rate
significantly), vs identical liquid 0 to 5 C. This phenomenon also remains
unexplained
(some possible answers are considered below).
1 S Of note is the last animal given machine ventilation (FIG. 1 I d) which
also was given
adequate gas ventilation, since the C02 difficulty in warm MMLV was known.
This
animal had no C02 buildup despite full rates of infusion (in fact. at rates
slightly higher
than standard-63 mL/kg/min). This data suggests that the C02 problem can be
solved, at
least in the sort term, by increasing gas ventilation with MMLV-a luxury not
available
with TLV. Calculations of minute volumes using C02SM0 data suggest that these
animals
had been under ventilated with gas by at least 50% with MMLV, with no
compensation, but
had withstood this without problem in the cold. On analysis of data from these
experiments, it was found that at these very fastest PFC infusion rates,
pressure limitations
were forcing a large decrease in volume of inspired gas breaths, and thus gas
ventilation
was not proceeding optimally. When the protocol was altered to allow small
infusions (8
mL/kg/min) and slightly more time for normal gas ventilation (with 10 seconds
given for
infusion and 5 extra seconds for full gas ventilation allowed after each 5-
second suction, for
a 50% infusion duty cycle), C02 was maintained at normal levels during lavage
with 62.5
mL/kg/min body temperature PFC, over 18 minutes (see FIG.IIb). These are
essentially
identical total PFC total ventilation rates (62 mL/kg/min at 50% of time = 31
mL/kg/min)
rates as used for the other animals in this series. These data are evidence
that MMLV is
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capable of achieving adequate C02 removal under many circumstances, in sharp
contrast to
the results obtained in TLV over comparable time courses
FIG.12 shows preliminary data explaining why cold MMLV experiments have less
problems with C02 elimination. FIG.12 shows combined EtC02 and VC02 data
during
cooling (Example 6) by 5°C using MMLV. Preliminary examination of this
data reveals a
profound decrease in both EtC02 and VC02 which would seem to indicate
decreased
C02 elimination. However, these data are misleading since they reflect only
the C02
which is present in the expired gas from mechanical ventilations. An
additional amount of
C02 is removed in the PFC which is loaded and unloaded from the lungs, and
this
volume of C02 is not measured by the capnograph in the gas path since it is
dissolved in
the PFC liquid and removed with it.
Evidence that the missing volume of C02 is not retained in the animal during
MMLV is seen at the conclusion of the experiment where both EtC02 and VC02
return
to near baseline levels and continue to remain at normal levels for over 1.5
hours without
significant rebound. An arterial blood gas drawn 3 minutes after the procedure
revealed a
temperature-corrected pa02 of 535.9 mmHg, a paC02 of 35.8 mmHg and a pH of
7.380
indicating an absence of hypercarbia and associated respiratory acidosis at
the end of the
experiment. This was true even though the VC02 had remained close to baseline.
The
temperature-corrected paC02 prior to the start of MMLV was 45.1 mmHg.
Quantitative Discussion of C02 Removal During Experiment Example 6
It is apparent from the data values for minute ventilation values using CO,SMO
data, that that very significant amounts of gas are being delivered to the
lungs throughout
the liquid loading and unloading cycles. Indeed, Vti is 70% of baseline. This
implies that
significant gaseous alveolar ventilation is continuing throughout liquid
loading and
unloading cycles thus greatly facilitating gas exchange and allowing it to
occur more or
less independently of liquid "breaths."
Total minute gas ventilation as noted above is reduced to 70 % of baseline and
yet
C02 removal in ventilatory gas falls to ~40% of baseline. Since C02 is not
building up
in the animal this implies that ~60% of C02 production is being removed in the
suctioned
PFC. Therefore a 30% reduction in gas ventilation is more than compensated for
by PFC
ventilation. These rates of C02 removal via PFC are considerably greater than
have been
previously reported in the literature for TLV at these minute volumes of PFC.
The
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improvement of C02 removal via PFC ventilation in MMLV is thus most likely due
to
good C02 equilibration in most of the volume of infused liquid, with a
consequently
small effective dead space. This results from "convective" mixing of PFC by
gas stirring,
in MMLV.
A consequence of minimization of the effective dead space (diffusion dead
space
plus anatomical dead space) is that smaller minute volumes of PFC may be used
in
ventilation and heat exchange without the loss of efficiency which occurs as a
result of
dead space volumes which are large fractions of liquid minute volumes. For
instance, in
Example 6 ~60% of C02 removal is carried out via PFC ventilation. This is
approximately 50% of the volume of PFC normally used in TLV in dogs. In TLV,
minute volumes of liquid may not be arbitrarily decreased without loss of COz
removal
efficiency. However, in MMLV there is little loss of efficiency when this is
done.
As seen in FIG 12, the average minute C02 production (VC02) in this animal
with PEEP off, in the five minutes prior to beginning the experiment, was 97.2
mL/min
1 S at standard temperature and pressure (STP). (Note that the one minute
averaging at the
beginning and end of the experiment result in false values which have been
dropped from
calculations). During the experiment (see Example 6), average minute C02
production
(as measured in expired gas) dropped to 38.3 mL/min (STP). After the end of
the
experiment, average minute VC02 increased again to 88 mL/min (STP) in the 7
minutes
before PEEP was re-instituted (some of this decrease is due to lowered
metabolic rate
after cooling). An average C02 production of 93 mL/min before and after PFC
cooling
gives a deficit of 93 - 38 ~ 55 mL C02 (STP)/min. In this 19.8 kg animal, this
deficit
amounts to 2.8 mL of C02 (STP)/kg/min, which is carried away dissolved in the
PFC.
According to 3M Company literature (3M Product Information Notebook), the
proprietary PFC used in this experiment (3M "FC-75" =
prefluorobutyltetrahydrofuran)
has a C02 carrying capacity of 2 mL C02 (STP) per mL of PFC at 760 Torr and
25°C
average PFC suction temp, see FIG.9). At the arterial pC02 measured just after
completion of the experiment (PaC02 = 35.8 Torr), carrying capacity of the PFC
at the
suction temperature would be estimated by Henry's law to be approximately 2.1
mL C02
(STP)/mL PFC x 35.8 = 760 = 0.098 mL C02 (STP)/mL PFC. At the average infusion
removal PFC cycle rates in this experiment (27 mL/kg/min), this would result
in a
theoretical C02 removal capacity of 27 mL PFC/kglmin x 0.098 mL C02 (STP)/mL
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PFC = 2.65 mL C02 (STP)/kg/min. This is very close to the amount of C02
deficit
measured: 2.8 mL C02 (STP)/kg/min. At 0°C the carrying capacity is
calculated to
increase by 6%, which is enough to raise the theoretical C02 carrying capacity
of the
PFC in Example 6 to the measured 2.8 mL C02 (STP)/kg/min. Supersaturation of
warm
PFC with gas dissolved at colder temperatures may also be operating to raise
carrying
capacity. Another possibility is that a small amount (6 % ) of the C02 evolved
which is
not recorded by the capnograph, is not dissolved in PFC but is instead removed
in the
suctioned gas which cannot be evaluated by the capnograph. In any case, it is
apparent
from this example that MMLV results in very close to theoretical saturation of
cold PFC
infusions with C02 - further indication that physiologic and diffusion dead
space with
MMLV have been greatly reduced when compared with TLV.
FIG. 13 shows VC02 measurements of another animal given cold PFC, again with
the suggestion that C02 production does not go up much (if at all) due to 18
minutes of
cold lung lavage. The VC02 increase with shivering and bringing body
temperature up
after the procedure, however, may be clearly seen. This is an example of a
hypermetabolic state which MMLV should be able to address in clinical
situations.
Other Physiologic Parameters
Electrolytes and lactate levels remained unchanged prior to, during, and after
the
experiment in Example 6. Serum lactate was 2.0 mM prior to the start of MMLV,
1.6
mM 3 minutes after MMLV and 1.3 mM 3 hours following the procedure.
The hemodynamic effects of MMLV are illustrated for Example 6. FIG.14 shows
the relationship between aortic temperature and heart rate. The temperature
wave form
shown in this graph is a surrogate for the volume of liquid in the lung. It
should be noted
that of the 22 dogs subjected to ultra cold intrapulmonary cooling with PFC,
no animal
developed any arrhythmia other than bradycardia associated with cooling, and
this was
transient, even at core drops of -12°C.
In FIG. 14 the peaks in aortic blood temperature indicate the point where most
of
the liquid is unloaded from the lungs. Similarly, the nadirs in aortic blood
temperature
are indicative of the presence of the maximum volume of cold PFC in the lungs.
As can
be seen from this graph, maximal PFC loading (and the nadir of aortic blood
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temperature) are associated with a modest decrease in heart rate. Initially it
was believed
that this decrease in heart rate was due to coronary blood cooling resulting
in a decrease
in myocardial metabolism with a associated decline in heart rate. This may in
fact explain
some of the decrease in heart rate, however, as can be seen from FIG.15, there
exists a
S better correlation between decrease in heart rate is the increase in central
venous pressure
(and thus intra-thoracic pressure) associated with PFC loading.
The more probable mechanism of this decrease in heart rate is decreased
preload as
a result of reduced venous return secondary to increased infra-thoracic
pressure peaking
at the end of PFC infusion. Interestingly, heart rate rebounds above baseline
when liquid
is suctioned from the lungs and the venous return to the heart normalizes.
This is
analogous to the reflexive rebound in heart rate after the release of pressure
in the
Valsalva maneuver. This phenomenon occurs since intrapleural (infra-thoracic)
pressure
has a profound effect on venous blood return to the heart and thus on preload.
Normal infra-thoracic pressure is -4 mmHg. An increase in infra-thoracic
pressure
to +2 mmHg requires a 6 mmHg increase in right atrial pressure (RAP). In the
absence
of compensatory circulatory reflexes, a rapid increase in RAP to 7 mmHg
decreases
venous blood return to zero. Even a modest rise in RAP as a result of PFC
loading causes
a drastic decrease in venous return because the systemic circulation is a very
flexible
compartment with the result that any increase in RAP causes blood to
accumulate in this
compartment and not return to the heart.
Lack of adequate preload (decrease in venous return to the heart) decreases
cardiac
output which in turn decreases arterial pressure. As can be seen from FIG.16,
the effect
of increased CVP on mean arterial pressure (MAP) is even more profound than
the effect
of PFC loading on heart rate. Maximum intrapulmonary PFC load is associated
with a
temporary drop in MAP of 15 to 25 mmHg. This decrease in MAP is directly
related to
increased infra-thoracic CVP and thus infra-thoracic pressure. As shown in
FIG.17, CVP
increases from an average of 2.0 mmHg to 3-3.5 mmHg at the maximum
intrapulmonary
PFC load of 50 ml/kg.
The relationships discussed above may also been seen in dogs loaded with body
temperature PFC (FIGs 16a and 17a), so there is no question that the
bradycardia and
acceleration associated with suction in these animals is an artifact primarily
of thoraic
pressure changes, not sudden blood hypothermia.
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FIG.18 shows the relationship between PFC load in the lungs (as indicated by
PFC
infusion/suction temperature curves) and ventilator gas pressure in Example 6.
In this
graph, PIP (peak inspiratory pressure as indicated by peaks in the ventilator
pressure
graph) and MAP rise steeply and nonlinearly as maximal PFC loading is
achieved. In this
experiment, PIPS are at 40 cm H20 by the time the last gas breath is delivered
with the
lungs maximally loaded with PFC. As can be seen from this data, the most
effective gas
breaths are being delivered at pressures of 25-27 cm H20 or less. PIPs of 40
cm H20 are
associated with rapid development of baro- and volu-trauma in the dog lung.
The
development of a more sophisticated apparatus for delivering MMLV will allow
for
lower PIPs and mean airway pressures with consequent reduction in lung injury.
Effect of MMLV on Respiratory Parameters and Some Caveats About the C02SM0
Total alveolar minute ventilation decreases proportionally more than total gas
minute volume, as computed by the C02SM0 device, during MMLV. This occurs
because the C02SM0 computes minute alveolar ventilation by using minute
ventilation
and VC02. Since 59% of the VC02 is being removed in the form of gas bubbles in
the
PFC liquid, the sensor of the C02SM0 cannot detect this volume of C02 and thus
the
values for gaseous C02 "alveolar minute volume" production (MV alv) total are
artificially low. However, calculation of the theoretical maximum amount of
C02 being
removed via PFC suction shows close agreement with the estimated VC02 of the
animal
during MMLV. Other caveats that should be considered when using the C02SM0 in
MMLV are:
a) The presence of PFC vapor in the respiratory gas being evaluated by the
C02SM0 results in falsely high VC02 and EtC02 readings on the order to of 6 to
13
(ca. 2-5 ml/min or 2-5 mmHg). Similarly, flow readings are modestly effected
(5 to 10% )
by the increased density and viscosity of ventilating gas loaded with PFC
vapor.
b) Positive VTi spontaneous values during MMLV do not reflect spontaneous
respiration since the animal is anesthetized (Level III, Plane IV) and
paralyzed throughout
the experiment. The C02SM0 is designed to report negative pressure airflow
down the
ET tube as spontaneous inspiration. In fact, the large values for VTi are an
artifact of
suctioning PFC and gas from the lungs of the animal and the ventilator circuit
at the
completion of liquid unloading. Arguably, these values constitute some
information about
the total volume of gas/liquid removed from the animal and ventilator circuit
during a
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liquid unloading cycle (since all liquid/gas suctioned from the lungs must be
replaced by
gas flowing through the C02SM0 over multiple cycles of suction) and thus may
be of
some use.
c) Wetting the capnographic window with PFC results in false high readings of
both EtC02 and VC02 until the heater in the capnograph vaporizes the PFC. The
dual
beam infrared technology used in the capnographic sensor does not tolerate
contamination
with liquid, whether it is PFC or water.
d) Similarly, the C02SM0 is designed to measure respiratory parameters using
only gas ventilation. The system operates by measuring pC02 and the absolute
and
differential pressures of the ventilating gas. Liquid contact, including PFC,
renders the
system inoperative.
The above caveats on the limitations of the C02SM0 data notwithstanding, it
should be noted that following completion of the experiment and suctioning of
PFC from
the lungs to FRC, the total alveolar minute volume returns to baseline levels
within 1-2
minutes where it remains for more than 2 hours until recovery from anesthesia
and the
start of weaning from conventional mechanical ventilation.
Further, despite limitations on its accuracy, the C02SM0 provides valuable
data
on the respiratory mechanics and physiology of MMLV. Recent changes to the
C02SM0
software allow for continuous real-time acquisition of data, as opposed to 1-
minute
trends. Use of this enhanced capability should allow for considerable progress
in
optimizing the algorithms used for MMLV.
Progress in Preventing Baro-trauma
FIGs 19 and 20 show 2 cold PFC lavage trials from 10 and 10a, in terms of
maximal
airway pressures, measured at the ET cuff. They differ essentially only by
infusion
volume (not rate) and machine control (machine control of valuing, and a newer
ventilator which is more sophisticated). While control of pressure is good
with practice in
hand-controlled mode, it is clear that the machine-controlled mode with
infusions of 8 mL
vs. 16 mL and better control over pressure parameters, did the best at
controlling
pressures in both suction and lavage.
Histopathological Evaluation
Blinded histopathological examination was performed by Ronald E. Gordon,
Ph.D., Director of the Electron Microscopy Core Facility at Mt. Sinai School
of
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Medicine, New York, NY. Tissues were evaluated grossly and by light microscopy
using
Hematoxylin and Eosin staining. All animals evaluated histopathologically were
subjected
to manual MMLV using the perfluorocarbon FC-75 per the protocols described
herein.
(No histopathology for machine controlled MMLV dogs is presently available,
but these
animals appear to do better clinically).
Example 2, FIGs. 3 and 4: The lungs appeared grossly normal. On light
microscopy there were occasional small areas of focal edema and hyper-
cellularity in the
bronchioles. The heart was grossly and microscopically normal.
Full MMLV Examples:
I 0 MMLV 1: The lungs looked good overall with normal gross and microscopic
appearance. Large airways, smaller airways, bronchioles, alveolar ducts,
alveoli and
blood vessels appeared normal. The gross and microscopic appearance of the
heart was
normal.
MMLV2: The lungs appeared grossly normal. The large and small airways were
I S intact, however, the vessels appeared dilated. The bronchioles and gas
exchange
compartments exhibited some interstitial edema with some RBC extravasation.
Vessels
were apparently acutely disrupted. There was no evidence of inflammatory
infiltrate.
NOTE: it is the pathologist's impression that this is an acute injury probably
secondary to
excessive perfusion pressure during fixative perfusion due to the lack of
inflammatory
20 changes unless this animal was sacrificed acutely. (In fact this animal was
clinically well
with normal ABGs at the time of sacrifice which was 2 weeks post MMLV). The
heart
was normal on both gross and microscopic examination.
MMLV3: The lungs appeared normal in all respects except for a few scattered
focal bronchiolar inflammatory infiltrates. The heart was normal both grossly
and on
25 microscopic examination.
MMLV4: The lungs and heart appeared completely normal.
As is evident from these histopathological findings, MMLV as practiced even in
its
currently suboptimum form results in only minor histological injury. As the
preceding
clinical and laboratory data make clear, MMLV is consistent with both good
post-
30 procedure gas exchange, ventilatory mechanics, and long term survival with
no clinically
apparent sequelae.
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Optimization of the Current Protocol for MMLV
In much C02SM0 data it can be seen that with hand-controlled relatively large
lavage loads, 1-2 gas breaths that are delivered before PFC loading of the
lung is at
maximal levels, are too often delivered at a PIP of 40 cm H20 or greater. PIPs
of 40 cm
H20 in dogs are associated with baro- and volu-trauma as well as pulmonary
edema due
to increased microvascular permeability. As can be seen from measured airway
pressures
vs. PFC load in the lungs, as shown in FIG.18, even in the worst of our
manually
controlled series, PIPS during MMLV averaged only 35 cm H20. Thus, no animals
were
permanently damaged by MMLV alone. Nevertheless it is still possible that
damage was
done which would not be acceptable clinically in certain situations. Thus,
optimization of
the protocol for certain clinical applications, can include one or more of the
following:
a) Substitute a sophisticated (feedback responsive) pressure cycled ventilator
for the
volume cycled, pressure limited ventilator (Puritan Bennett MA-1) which was
used to do
much of the manual work and used for most of these experiments. FIG 31 shows
not
only machine controlled ventilation, but the ventilator is a Servo-9000.
b) Limit PIP to no higher than 25 to 27 cm H20 and mean airway pressure to no
greater than 10 cm H20.
c) Eliminate large negative intrapulmonary pressures during PFC suctioning and
halt suctioning at a negative airway pressure of 2 to 4 cm H20. The rate of
heat exchange
might be further optimized by discontinuing suction at a PIP of +2 to +5 cm
H20 at
which time almost all of the available PFC would be unloaded from the lung.
This would
not only avoid any possibility of trauma from negative intrapulmonary
pressures, but
would optimize heat exchange by increasing the frequency of liquid
loading/unloading
cycles per unit of time.
CONCLUSION
This document enumerates multiple uses for gas/PFC-based liquid ventilation
(MMLV) as a mode of efficient heat exchange due to the mixing effect of gas
breaths on
heat exchange medium in the small airways throughout the lung, including in
the
peripheral lung (gas bubble-induced small-scale PFC non-thermal "convection").
This
can be used in the entire range of possible liquid infusion temperatures (-
10°C to 43°C),
since prior art in PFC heat exchange relies on non augmented TLV with no gas
bubbles
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present (Shaffer, 1994, US Patent # 5,335,650, and Vaseen VA, 1980, US Patent
#
4,323,665).
There are other means of deliberately inducing small airway liquid mixing in
PFC
for purposes of heat exchange, as being essentially the same as the gas bubble
method
induced herein. These include but are not limited to:
A. The use of High Frequency [Oscillating] Ventilation (HFOV, or HFV) as a
modality to induce small airway fluid mixing in PFC, for purposes of heat
exchange.
This will be true for HFOV when used with either TLV and MMLV (with and
without
gas, since HFOV is expected to cause nonthermal "convective" mixing and gas
transport,
even in TLV). HFOV has been used experimentally to obtain good C02 "dead
space"
diffusion with small tidal volumes (i.e., equal to the anatomical dead space)
in PLV with
PFCs. The suggestion and expectation that it may be used to augment heat
exchange with
PFC in both TLV and MMLV is novel.
B. The use of sweep-flow augmented PFC, introduced in prior art (Parker JC, US
patent # 5,706,830, 1996) as a means of reducing C02 "diffusion dead space" by
using
jets of PFC delivered via two canulae directed down the main bronchi. These
jets induce
dead space fluid mixing, but are introduced in prior art only as aids to C02
removal in
TLV. The heat exchanger in this patent serves only to buffer excess heat from
the
pumps, prevent hypothermia in the subject due to ventilation with room
temperature
liquid and prevent heat losses from the circuit tubing from inducing
hypothermia. This
technique is not used to deliberately to cool or warm the subject.
The proposal for the use of sweep-flow technique as an aid to heat transfer in
TLV,
PLV or MMLV, is novel and public disclosure of sweep-flow (increased
ventilatory
efficacy due to turbulent mixing of the anatomical dead space) independent of
J.C.
Parker, by one of us (M.Darwin) in April of 1996 predates the filing of J.C.
Parker's
patent # 5,706,830. This patent proposes to use a Y-configuration
endobronchial tube to
facilitate elimination of physiologic dead space.
C. The use of deliberate combinations of HFOV, sweep-flow, and MMLV to
induce small airway mixing of PFC based liquids or other suitable liquids, for
purposes
of heat exchange or gas exchange.
We have developed a device which uses gas ventilation specifically to
facilitat a
removal of PFC (or other liquid) liquid heat exchange and gas exchange media
and gas
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foam mixtures, from the lungs. These high rates of PFC or other liquid removal
are
necessary for the highest heat and gas exchange rates a non damaging
pressures.
In addition, MMLV as a technique can include a novel device which controls gas
and liquid ventilation separately, but in an interlocked way, in order to
maximize C02
S removal, heat removal and oxygen delivery.
These techniques have a number of uses, including but not limited to:
Delivering ultracold PFC infusion into the lungs, which we have found may be
used safely, without cold induced lung injury. We find that it can be used
safely in the
temperature range below 20°C (about -10°C to +20°C) for
PFC infusion of any kind
into the lungs. Prior art (Shaffer, 1994, US Patent # 5,33S,6S0} claimed the
use of TLV
for heat exchange only with fluids warmer than 20°C, but cooler than
body temperature.
We've introduced the novel idea of ultra low temperature (-10°C to
20°C) PFC infusion
for heat exchange, as used in any kind of liquid ventilation or liquid
breathing.
These techniques are not limited to the exact techniques described. It will be
noted
that MMLV is susceptible to improvements in many ways by anyone skilled in the
art,
without departing from the spirit of the discovery that non-thermal
"convective" mixing
and mass transport assistance is necessary for efficient central cooling via
PFC lung
lavage. Cooling rates of 20°C/hr are easily achieved by our methods,
but we note that
faster cooling rates are in principle easily achievable by maximizing
performance of many
features of our system of MMLV, and by use of features of systems which we
have
described above (HFOV, sweep-flow) when used in the novel way described (i.e.,
in
conjunction with PFC for heat exchange and/or gas exchange).
Increasingly, we've found that MMLV facilitates gas exchange more efficiently
for
the treatment of adult and neonatal respiratory distress syndromes, pulmonary
edema, and
2S other pulmonary insults (i.e., alveolar proteinosis, chronic
bronchiectasis, as well as
chemical and thermal insults to the lung) which result in V/Q mismatch as a
consequence
of sequestration of alveoli from gas exchange by means other than either TLV,
as taught
by Shaffer, 1994, US Patent S,33S,6S0, and Vaseen VA, 1980, US Patent #
4,323,665,
or PLV as taught by Schutt, EG 1996 in US Patent # S,S40,225. The use of
normothermic MMLV in these conditions has the added advantage of acting to
vigorously
lavage the large and small airways-- thus removing mucus, blood (secondary to
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hemoptysis or trauma), pulmonary transudate, and other harmful respiratory
secretions-
all far more efficiently than possible with either TLV or PLV.
MMLV can also be used for the therapeutic induction or reversal of
hypothermia,
including but not limited to: heatstroke, malignant hyperthermia,
hyperpyrexia, stroke,
head injury, post-ischemic insult, and febrile illnesses.
MMLV is useful for the companion animal and human cryopreservation patient
(cryonic suspension) and for other postmortem cooling or warming of humans for
the
purposes of organ preservation, organ or tissue recovery, resuscitation, or
facilitation of
treatment of trauma patients, exsanguinating injuries, or cardiac arrest.
MMLV is a therapeutic modality to improve gas exchange, reverse or induce
hypothermia, or maintain normothermia. This claim is understood to include but
not be
limited to shock as a result of sepsis, poisoning, chemical or thermal burns,
and trauma.
MMLV can be used for the purpose of increasing the efficacy of closed chest
CPR
by the mechanism of raising intra-thoracic pressure during the down stroke of
external
cardiac compression by synchronizing liquid loading with chest compression. A
corollary
of increased thoracic pressure during the downstroke of CPR is increased
cardiac output
as a result of decreasing lung compliance (due to liquid loading) thus
facilitating cardiac
output in CPR via the thoracic pump mechanism. It may be quite dangerous to do
CPR
in lungs fully loaded with liquid, which they need to be in TLV. However, we
have
found that boluses of oxygenated PFC can allow metabolic supply of oxygen for
many
minutes without ventilation at all. It may be that coordinated use of small
amounts of
PFC in CPR will allow less ventilation when chest compression needs to be
done, but still
allow for less chest compressibility, by reason of displacement of relatively
elastic gas in
the lungs.
Nitric oxide and nitric oxide donors administered via the breathing liquid or
the
breathing gas, or both, can be used to overcome cold-induced pulmonary
vasoconstriction
and thus facilitate gas and heat exchange. It is possible that pulmonary
vasoconstrictors
(nitric oxide antatonists) will be able to reverse the apparent V/Q mismatch
caused by
rapidly cycled normothermic PFC in normal lungs, without compromising system
circulation.
All, or a large fraction of the oxygen or other breathing gases in the gas
ventilation
component of MMLV can be replaced with helium in order to reduce ventilating
gas
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viscosity, thus allowing for lower peak airway pressures secondary to gas flow
in gas
ventilation. Helium is anticipated also to have unique properties in gas-
liquid foams: it
should form smaller gas bubbles, resulting in improved bubble -induced mixing.
The use
of helium will also facilitate nitrogen out-gassing from body water during re-
warming
from hypothermia. The use of 100% inspired helium gas is possible when used
with fully
oxygenated PFC delivered at appropriate infusion rates ( > 10 mL/kg/min).
-S 5-

Representative Drawing
A single figure which represents the drawing illustrating the invention.
Administrative Status

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Administrative Status

Title Date
Forecasted Issue Date Unavailable
(86) PCT Filing Date 1999-10-01
(87) PCT Publication Date 2000-04-06
(85) National Entry 2001-03-26
Examination Requested 2004-09-23
Dead Application 2008-10-02

Abandonment History

Abandonment Date Reason Reinstatement Date
2007-10-02 R29 - Failure to Respond
2007-10-02 R30(2) - Failure to Respond
2008-10-01 FAILURE TO PAY APPLICATION MAINTENANCE FEE

Payment History

Fee Type Anniversary Year Due Date Amount Paid Paid Date
Application Fee $300.00 2001-03-26
Maintenance Fee - Application - New Act 2 2001-10-01 $100.00 2001-03-26
Registration of a document - section 124 $100.00 2001-04-10
Maintenance Fee - Application - New Act 3 2002-10-01 $100.00 2002-09-27
Maintenance Fee - Application - New Act 4 2003-10-01 $100.00 2003-10-01
Request for Examination $800.00 2004-09-23
Maintenance Fee - Application - New Act 5 2004-10-01 $200.00 2004-09-24
Maintenance Fee - Application - New Act 6 2005-10-03 $200.00 2005-09-30
Maintenance Fee - Application - New Act 7 2006-10-02 $200.00 2006-09-29
Maintenance Fee - Application - New Act 8 2007-10-01 $200.00 2007-07-26
Owners on Record

Note: Records showing the ownership history in alphabetical order.

Current Owners on Record
CRITICAL CARE RESEARCH, INC.
Past Owners on Record
DARWIN, MICHAEL GREGORY
HARRIS, STEVEN BRADLEY
RUSSELL, SANDRA RENEE
Past Owners that do not appear in the "Owners on Record" listing will appear in other documentation within the application.
Documents

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Document
Description 
Date
(yyyy-mm-dd) 
Number of pages   Size of Image (KB) 
Representative Drawing 2001-06-26 1 10
Description 2001-03-26 55 3,194
Abstract 2001-03-26 1 64
Cover Page 2001-06-26 1 36
Claims 2001-03-26 3 116
Drawings 2001-03-26 29 775
Assignment 2001-03-26 4 122
Assignment 2001-04-10 3 114
PCT 2001-03-26 10 399
Prosecution-Amendment 2001-03-26 1 18
Correspondence 2001-06-12 1 20
Assignment 2002-06-25 1 47
Fees 2003-10-01 1 38
Prosecution-Amendment 2004-09-23 1 24
Fees 2004-09-24 1 40
Fees 2007-07-26 1 36
Fees 2002-09-27 1 42
Fees 2005-09-30 1 36
Fees 2006-09-29 1 36
Prosecution-Amendment 2007-04-02 3 126